WO2008033497A1 - Nanostructured smart gel for time release drug delivery - Google Patents

Nanostructured smart gel for time release drug delivery Download PDF

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Publication number
WO2008033497A1
WO2008033497A1 PCT/US2007/019992 US2007019992W WO2008033497A1 WO 2008033497 A1 WO2008033497 A1 WO 2008033497A1 US 2007019992 W US2007019992 W US 2007019992W WO 2008033497 A1 WO2008033497 A1 WO 2008033497A1
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thermo
lidocaine
water
reversible
gel
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PCT/US2007/019992
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French (fr)
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Benjamin Chu
Benjamin S. Hsiao
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The Research Foundation Of State University Of New York
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Publication of WO2008033497A1 publication Critical patent/WO2008033497A1/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
    • A61K9/5107—Excipients; Inactive ingredients
    • A61K9/513—Organic macromolecular compounds; Dendrimers
    • A61K9/5161—Polysaccharides, e.g. alginate, chitosan, cellulose derivatives; Cyclodextrin
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61K—PREPARATIONS FOR MEDICAL, DENTAL OR TOILETRY PURPOSES
    • A61K41/00—Medicinal preparations obtained by treating materials with wave energy or particle radiation ; Therapies using these preparations
    • A61K41/0028—Disruption, e.g. by heat or ultrasounds, sonophysical or sonochemical activation, e.g. thermosensitive or heat-sensitive liposomes, disruption of calculi with a medicinal preparation and ultrasounds
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61K—PREPARATIONS FOR MEDICAL, DENTAL OR TOILETRY PURPOSES
    • A61K47/00—Medicinal preparations characterised by the non-active ingredients used, e.g. carriers or inert additives; Targeting or modifying agents chemically bound to the active ingredient
    • A61K47/30—Macromolecular organic or inorganic compounds, e.g. inorganic polyphosphates
    • A61K47/34—Macromolecular compounds obtained otherwise than by reactions only involving carbon-to-carbon unsaturated bonds, e.g. polyesters, polyamino acids, polysiloxanes, polyphosphazines, copolymers of polyalkylene glycol or poloxamers
    • A—HUMAN NECESSITIES
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    • A61K47/50—Medicinal preparations characterised by the non-active ingredients used, e.g. carriers or inert additives; Targeting or modifying agents chemically bound to the active ingredient the non-active ingredient being chemically bound to the active ingredient, e.g. polymer-drug conjugates
    • A61K47/51—Medicinal preparations characterised by the non-active ingredients used, e.g. carriers or inert additives; Targeting or modifying agents chemically bound to the active ingredient the non-active ingredient being chemically bound to the active ingredient, e.g. polymer-drug conjugates the non-active ingredient being a modifying agent
    • A61K47/56—Medicinal preparations characterised by the non-active ingredients used, e.g. carriers or inert additives; Targeting or modifying agents chemically bound to the active ingredient the non-active ingredient being chemically bound to the active ingredient, e.g. polymer-drug conjugates the non-active ingredient being a modifying agent the modifying agent being an organic macromolecular compound, e.g. an oligomeric, polymeric or dendrimeric molecule
    • A61K47/59—Medicinal preparations characterised by the non-active ingredients used, e.g. carriers or inert additives; Targeting or modifying agents chemically bound to the active ingredient the non-active ingredient being chemically bound to the active ingredient, e.g. polymer-drug conjugates the non-active ingredient being a modifying agent the modifying agent being an organic macromolecular compound, e.g. an oligomeric, polymeric or dendrimeric molecule obtained otherwise than by reactions only involving carbon-to-carbon unsaturated bonds, e.g. polyureas or polyurethanes
    • A61K47/60—Medicinal preparations characterised by the non-active ingredients used, e.g. carriers or inert additives; Targeting or modifying agents chemically bound to the active ingredient the non-active ingredient being chemically bound to the active ingredient, e.g. polymer-drug conjugates the non-active ingredient being a modifying agent the modifying agent being an organic macromolecular compound, e.g. an oligomeric, polymeric or dendrimeric molecule obtained otherwise than by reactions only involving carbon-to-carbon unsaturated bonds, e.g. polyureas or polyurethanes the organic macromolecular compound being a polyoxyalkylene oligomer, polymer or dendrimer, e.g. PEG, PPG, PEO or polyglycerol
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    • A61K47/6921—Medicinal preparations characterised by the non-active ingredients used, e.g. carriers or inert additives; Targeting or modifying agents chemically bound to the active ingredient the non-active ingredient being chemically bound to the active ingredient, e.g. polymer-drug conjugates the conjugate being characterised by physical or galenical forms, e.g. emulsion, particle, inclusion complex, stent or kit the form being a particulate, a powder, an adsorbate, a bead or a sphere
    • A61K47/6927—Medicinal preparations characterised by the non-active ingredients used, e.g. carriers or inert additives; Targeting or modifying agents chemically bound to the active ingredient the non-active ingredient being chemically bound to the active ingredient, e.g. polymer-drug conjugates the conjugate being characterised by physical or galenical forms, e.g. emulsion, particle, inclusion complex, stent or kit the form being a particulate, a powder, an adsorbate, a bead or a sphere the form being a solid microparticle having no hollow or gas-filled cores
    • A61K47/6929—Medicinal preparations characterised by the non-active ingredients used, e.g. carriers or inert additives; Targeting or modifying agents chemically bound to the active ingredient the non-active ingredient being chemically bound to the active ingredient, e.g. polymer-drug conjugates the conjugate being characterised by physical or galenical forms, e.g. emulsion, particle, inclusion complex, stent or kit the form being a particulate, a powder, an adsorbate, a bead or a sphere the form being a solid microparticle having no hollow or gas-filled cores the form being a nanoparticle, e.g. an immuno-nanoparticle
    • A61K47/6931—Medicinal preparations characterised by the non-active ingredients used, e.g. carriers or inert additives; Targeting or modifying agents chemically bound to the active ingredient the non-active ingredient being chemically bound to the active ingredient, e.g. polymer-drug conjugates the conjugate being characterised by physical or galenical forms, e.g. emulsion, particle, inclusion complex, stent or kit the form being a particulate, a powder, an adsorbate, a bead or a sphere the form being a solid microparticle having no hollow or gas-filled cores the form being a nanoparticle, e.g. an immuno-nanoparticle the material constituting the nanoparticle being a polymer
    • A61K47/6935—Medicinal preparations characterised by the non-active ingredients used, e.g. carriers or inert additives; Targeting or modifying agents chemically bound to the active ingredient the non-active ingredient being chemically bound to the active ingredient, e.g. polymer-drug conjugates the conjugate being characterised by physical or galenical forms, e.g. emulsion, particle, inclusion complex, stent or kit the form being a particulate, a powder, an adsorbate, a bead or a sphere the form being a solid microparticle having no hollow or gas-filled cores the form being a nanoparticle, e.g. an immuno-nanoparticle the material constituting the nanoparticle being a polymer the polymer being obtained otherwise than by reactions involving carbon to carbon unsaturated bonds, e.g. polyesters, polyamides or polyglycerol
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    • A61K47/50—Medicinal preparations characterised by the non-active ingredients used, e.g. carriers or inert additives; Targeting or modifying agents chemically bound to the active ingredient the non-active ingredient being chemically bound to the active ingredient, e.g. polymer-drug conjugates
    • A61K47/69—Medicinal preparations characterised by the non-active ingredients used, e.g. carriers or inert additives; Targeting or modifying agents chemically bound to the active ingredient the non-active ingredient being chemically bound to the active ingredient, e.g. polymer-drug conjugates the conjugate being characterised by physical or galenical forms, e.g. emulsion, particle, inclusion complex, stent or kit
    • A61K47/6921—Medicinal preparations characterised by the non-active ingredients used, e.g. carriers or inert additives; Targeting or modifying agents chemically bound to the active ingredient the non-active ingredient being chemically bound to the active ingredient, e.g. polymer-drug conjugates the conjugate being characterised by physical or galenical forms, e.g. emulsion, particle, inclusion complex, stent or kit the form being a particulate, a powder, an adsorbate, a bead or a sphere
    • A61K47/6927—Medicinal preparations characterised by the non-active ingredients used, e.g. carriers or inert additives; Targeting or modifying agents chemically bound to the active ingredient the non-active ingredient being chemically bound to the active ingredient, e.g. polymer-drug conjugates the conjugate being characterised by physical or galenical forms, e.g. emulsion, particle, inclusion complex, stent or kit the form being a particulate, a powder, an adsorbate, a bead or a sphere the form being a solid microparticle having no hollow or gas-filled cores
    • A61K47/6929—Medicinal preparations characterised by the non-active ingredients used, e.g. carriers or inert additives; Targeting or modifying agents chemically bound to the active ingredient the non-active ingredient being chemically bound to the active ingredient, e.g. polymer-drug conjugates the conjugate being characterised by physical or galenical forms, e.g. emulsion, particle, inclusion complex, stent or kit the form being a particulate, a powder, an adsorbate, a bead or a sphere the form being a solid microparticle having no hollow or gas-filled cores the form being a nanoparticle, e.g. an immuno-nanoparticle
    • A61K47/6931—Medicinal preparations characterised by the non-active ingredients used, e.g. carriers or inert additives; Targeting or modifying agents chemically bound to the active ingredient the non-active ingredient being chemically bound to the active ingredient, e.g. polymer-drug conjugates the conjugate being characterised by physical or galenical forms, e.g. emulsion, particle, inclusion complex, stent or kit the form being a particulate, a powder, an adsorbate, a bead or a sphere the form being a solid microparticle having no hollow or gas-filled cores the form being a nanoparticle, e.g. an immuno-nanoparticle the material constituting the nanoparticle being a polymer
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Definitions

  • the present invention is directed to a unique class of nanostructured smart gels for site- specific pain management.
  • Chronic pain has been traditionally defined as pain persisting for at least 3-6 months but can now be defined as pain that extends beyond the period of tissue healing and/or with low levels of identified pathology that are insufficient to explain the presence and/or extent of the pain.
  • the persistence of chronic pain can affect many different aspects of personal life, ranging from the physiological to societal and financial changes. Chronic pain afflicts 24% of all Americans, most prominently affecting women who are over 50 years of age. Chronic pain patients typically average 12.9 trips to a General Practitioner per year - as opposed to the average 4.2 trips per year for the general community. Overall, the economic burden of the lost productivity due to chronic pain totals approximately $86.2 billion per year for the United States, a figure sure to exponentially rise as the "Baby Boomer" generation ages and no definitive therapeutic treatments have been discovered.
  • Lidocaine is one of the most widely used anesthetic drugs today, especially during surgery and dental procedures. Lidocaine was developed by the Swedish scientist Nils L ⁇ fgren in 1943. An additional use of lidocaine is as an anti-arrhythmic agent. The drug works by inhibiting the stimulants needed to initiate neuronal impulses to the brain, resulting in the loss of pain. As a Docket No. 788-74 (R7921) topical drug, lidocaine has a relatively short half-life of only 1.5-2 hours in an intravenous injection because it is quickly metabolized by the liver (due to the presence of an amide group).
  • lidocaine Even though the time frame in which lidocaine works is extremely short, it is commonly used as the local anesthetic of choice among professionals due to its hypoallergenic quality. Allergic reactions to lidocaine are extremely rare, and if they do occur, it is usually due to one of the preservatives found in the dose vials.
  • This present invention uses a thermally responsive carrier for controlled delivery of a known and desired amount of lidocaine to an injured site with pre-designed release profile over a flexible time period from many hours to 10-15 days.
  • the thermo-reversible sol-gel carrier which is a liquid at room temperature and a gel at body temperature, allows versatile deployment methods to be implemented during medical procedures.
  • An aspect of the present invention is to substantially solve at least the above problems and/or disadvantages and to provide at least the advantages below.
  • the present invention provides a way to provide the pain relief attributes of lidocaine as well as other therapeutic agents, including as prilocaine, bupivacaine, ropivacaine molecules, analogs and mixtures thereof, to a patient for a sustained period of time without affecting the attributes of the drug.
  • the invention utilizes smart gels as a drug delivery system.
  • the therapeutic agent(s) can be added directly to the smart gel itself for faster release, as well as to particles incorporated into the smart gel for sustained release.
  • the (sub-)micron size particles of the invention contain at least one therapeutic agent, such as lidocaine, prilocaine, bupivacaine, ropivacaine molecules, their analogs or their mixtures.
  • the particles are based on alginate/chitosan polyelectrolyte-complex (PC), chitosan polyelectrolyte-complex (PC) or mixtures.
  • the polyelectrolyte complex.alginate/chitosan forms one ion pair; but there are many others as well that Docket No.
  • the particle size ranges from about 10 am to about 1 micon; the alginate/chitosan ratio can be from 100/0 to 0/100.
  • the present invention provides a time-release system that allows the desired prolonged drug delivery in pain management.
  • he smart gel of the present invention is a a thermo-reversible gel composition.
  • Said gel can contain therapeutic agent(s) added directly thereto for faster release, as well as to particles incorporated into the smart gel for sustained release.
  • said particles contain therapeutic agents, such as lidocaine prilocaine, bupivacaine, ropivacaine molecules, their analogs or their mixtures; and alginate polyelectrolyte- complex (PC), chitosan polyelectrolyte-complex (PC), and mixtures thereof.
  • thermo- reversibility is obtained by changing the temperature of the solution and breaking the hydrogen bonds that are formed at a lower temperature and thereby causing the sol to gel and release the therapeutic agent is slowed down by the polymer network, whose mesh size can partially be controlled by the copolymer concentration.
  • the diffusion of therapeutic agent through the polymer network is much slower when compared with that in a polymer solution and thereby in slowing down the time release of the therapeutic agent.
  • thermo- reversible Pluronic hydrogel having a series of (E) and (P) triblock copolymers with a general formula of E x PyE x , wherein E is a hydrophilic molecule and P is a molecule that is capable of temperature-dependent hydrogen bonding with water (near the body temperature) and therefore is hydrophilic at low temperatures, but becomes hydrophobic at higher temperatures due to the break down of hydrogen bonds between the P blocks and water.
  • the value of x can range from about 1 to about 300 and the value of y can range from about 1 to about 300.
  • Still another embodiment of the present invention is directed to a smart gel comprising a thermo-reversible Pluronic hydrogel comprising a series of (P) and (E) triblock copolymers with the general formula Of P x EyP x .
  • the value of x can range from about 1 to about 300 and the value of y can range from about 1 to about 300.
  • the switching of roles between E and P, i.e., from E x PyE x to P x EyP x changes the morphology of the polymer network formed.
  • thermo-reversible Pluronic hydrogel compositions described above can be blended together to produce a single composition having combined properties.
  • the thermo-reversible Pluronic hydrogel blend comprising a series of (E) and (P) triblock copolymers with the general formula of E x PyE x , wherein E is a hydrophilic molecule and P is a molecule that is capable of a stronger temperature dependent hydrogen bonding with water and therefore is hydrophilic at low temperatures, but becomes hydrophobic at higher temperatures due to the break down of hydrogen bonds between the P blocks and water.
  • the value of x can range from about 1 to about 300 and the value of y can range from about 1 to about 300.
  • the composition also includes a series of (P) and (E) triblock copolymers with the general formula of P x E y P x , wherein E is a hydrophilic molecule and P is a molecule that is capable of a stronger temperature dependent hydrogen bonding with water (near body temperature) and therefore is hydrophilic at low temperatures, but becomes hydrophobic at higher temperatures due to the break down of hydrogen bonds between the P blocks and water having a value of x that ranges from about 1 to about 300 and a value of y that ranges from about 1 to about 300.
  • the composition contains at least one therapeutic agent, such as lidocaine, and results in a time-released composition. Docket No. 788-74 (R7921) BRIEF DESCRIPTION OF THE FIGURES
  • Figure 1 is the structure of lidocaine hydrochloride
  • Figure 2 is the Sol-gel phase diagram of aqueous 30% (w/v) F87/F127 mixture solutions
  • Figure 3a shows typical static light scattering results for the determination of cmt of F87 solutions at various concentrations
  • Figure 3b graphically shows the dependence of critical micelle temperature on polymer solution concentration
  • Figure 4 illustrates a typical SAXS intensity profile of 30% (w/v) F87 aqueous solution at 42 0 C;
  • Figure 5 shows the change of the lattice constant with the F87 content in the mixed solution
  • Figure 6 schematically shows representations of F87 and F127 triblock copolymer micelles in the cubic structures
  • Figure 7 graphically illustrates the temperature dependence of the zero-shear rate viscosity of 30% (w/v) F87/F127 mixed solution at a weight ratio of 1 :2 in aqueous solution;
  • Figure 8 graphically illustrates the temperature dependence of the zero-shear rate viscosity of 30% (w/v) F87/F127 mixed solutions (weight ratio of 1:2) at different concentrations of lidocaine;
  • FIG. 9 graphically illustrates the amount of lidocaine delivery, representing the fastest diffusion limit of lidocaine release
  • Figure 10 graphically illustrates the typical release profiles from the solution containing 2 wt% chitosan (10 mL), 4.4 wt% alginate (5 mL) and mineral oil (10 mL); Docket No. 788-74 (R7921)
  • Figure 11 schematically represents the general behavior of the pluronics of the present invention.
  • Figure 12 shows the phase diagram at 37 0 C for copolymer blends of F127 and F87 in water.
  • Figure 13 schematically shows a confined impinging jets (CU) mixer.
  • lidocaine hydrochloride (lidosalt), as shown in figure 1 and lidocaine base (lidobase).
  • Lidocaine hydrochloride with the formula Of Ci 4 H 22 N 2 O • HCl is the anesthetically active form and is soluble in water
  • lidobase with the formula Of Ci 4 H 22 N 2 O is not soluble in water and anesthetically much less active.
  • Lidocaine hydrochloride is predominately marketed by AstraZeneca as Xylocaine®.
  • Lidocaine hydrochloride has a very interesting structure in water.
  • the hydrate microcrystal theory of anesthesia, advanced by Linus Pauling is closely related to the "iceberg” theory of ionic solutions and hydration of proteins.
  • the ordered arrangement of water molecules around solute ions and protein side chains is considered as part of the clathrate structure.
  • Later studies have shown that local anesthetics form a hydrogen bonded complex with a receptor in the membrane.
  • Lidocaine, a sodium ion channel blocker or nerve block is a local anesthetic.
  • a lidocaine cation can donate two protons and accept one.
  • the crystal structure of lidocaine indicates that adjacent chains are held Docket No.
  • Infrared spectra of lidocaine cation with different anions indicate strong interactions between nitrogen and hydrogen and a small anion associated with an intense force field such as chloride.
  • an intense force field such as chloride.
  • the N + -H stretching frequency is higher than that with chloride ion.
  • the neutral form of lidocaine is about five times less effective when compared with the cationic form.
  • the choice of an appropriate anion may also influence the rate of drug delivery from the gel.
  • lidocaine because of its charge, is located in the outer aqueous region of the gel rather than in the hydrophobic interior of the micelles. This supposition has not been verified experimentally. Moreover, crystal data suggest that it is possible for lidocaine cations to enter the micelle by hiding its charge with the formation of 'endless' chains in the presence of chloride ions and hydrogen bonding with water. This observation necessitates a further need for data from laser light scattering as well as small angle and wide angle X-ray scattering.
  • lidocaine hydrochloride reported herein indicate negative differential resistance at some potential, suggesting possible tunneling of electrons Docket No. 788-74 (R7921) through the self assembled pathway or through the ordered arrangement of water molecules in between the chains near the electrode double layer.
  • Hydrogels are often used in pharmaceutical and biotechnological industries because of its low toxicity, little to no immune response and unique temperature-dependent viscosity. Hydrogels are the closest synthetic materials to emulate living tissue due to its ability to imbibe large amounts of water as well as its soft consistency. The amount of water that a hydrogel can take in is limited by the force of mixing as well as the retractive force of the polymer. Therefore, there is an upper limit as to how much of the hydrogels will, in fact, be dissolved. These factors greatly influence its low immune response.
  • Pluronic hydrogels are a series of ethylene oxide (E) and propylene oxide (P) triblock copolymers with the general formula of E x PyE x , which can be considered as amphiphilic non-ionic surfactants.
  • E ethylene oxide
  • P propylene oxide
  • their aqueous solutions exhibit liquid-like properties at low temperatures and gel-like properties at higher temperatures because the E blocks are consistently hydrophilic, which allows for the copolymer to be dissolved in water.
  • the middle P block is hydrophilic at low temperatures, but hydrophobic at higher temperatures due to the break down of hydrogen bonds between the P blocks and water.
  • This distinctive property of thermally responsive Pluronic hydrogels makes it easy to administer to patients.
  • the solution can be given through injection at a low temperature, and once the liquid enters the body, it would immediately gel.
  • the physically cross-linked gel would stay essentially in place right at the zone of treatment without having to worry about the drug impacting other areas of the body.
  • the thermally responsive gelation behavior of Pluronic copolymers changes with the E/P ratio and the total chain length.
  • BCC body centered cube
  • FCC face centered cube
  • PSCs polyelectrolyte-surfactant complexes
  • hydrophilic domains with surfactant head groups and polyelectrolvte charges
  • hydrophobic regions with surfactant tails and hydrophobic parts of the polyelectrolvte.
  • the two most important driving forces for the self- assembly of surfactant molecules in PSCs are: (1) electrostatic interactions between the charged components and (2) hydrophobic interactions between the polymer backbone and the alkyl chains of the surfactant.
  • PCs polyelectrolvte complexes
  • the oppositely charged polyelectrolytes are neutralized via electrostatic interactions, forming ion pairs and rendering hydrophobic regions, such as in the neutralization of negatively charged alginate and positively charged chitosan, in which both components are degradable in water (alginates are produced by brown seaweeds - phaeophyceae, mainly laminaria; and chitosan can come from a bacterial polysaccharide).
  • alginates are produced by brown seaweeds - phaeophyceae, mainly laminaria; and chitosan can come from a bacterial polysaccharide).
  • the drug molecules can be entrapped in the complexes for a prolonged period of time without diffusing out of the complex domain, as has been demonstrated by our preliminary experiments. Docket No. 788-74 (R7921)
  • Biodegradable microparticles have been routinely used in oral delivery systems and, even more often, in subcutaneously injected delivery systems because they can be administered to a variety of locations in vivo through a syringe needle.
  • the release rate of drugs from biodegradable microparticles can be controlled by a number of factors, such as polymer biodegradation kinetics, physicochemical properties of polymers and of drugs, the size and size distribution of microparticles.
  • drug release often exhibits an initial 'burst' phase during which a significant fraction (5-50%) of the encapsulated compound is released in a short time ( ⁇ 8 h).
  • microparticle size is an important determinant of drug release rate. Larger spheres generally release encapsulated compounds more slowly and over longer time periods. Thus, controlling microparticle size provides an effective and relatively simple pathway for control of drug release. Numerous studies have been conducted to determine the effects of microparticle size on drug release. Fabrication of biodegradable polymer microparticles with precise size and size distribution control provides a means for enhanced control of drug delivery rates.
  • microparticles with a relatively small size normally exhibit poor encapsulation efficiency and result in an undesirably rapid release of loaded drugs.
  • microparticles containing the desired drugs there are several techniques available for the production of microparticles containing the desired drugs, such as the emulsion-solvent evaporation/extraction method, spray drying and phase separation. Each technique has its own advantages and limitations. The choice of a particular technique depends on the attributes of the polymer and the drug, the site of drug action and the duration of the desired therapy. For example, the emulsion-solvent evaporation/extraction methods have been demonstrated to encapsulate both hydrophilic and hydrophobic drugs. However, the resulting particles often contain surfactant(s) that cannot be removed.
  • a microparticle system consisting of tailor-designed size distribution that can cover a broad range of drug release rates is more effective than a monodispersed microparticle system that can provide only a narrow range of drug release profile. Accordingly, a newly developed fabrication process was used to overcome the above concerns. The technique is termed Flash NanoPrecipitation using a confined impinging jets (CIJ) mixer, which is discussed in the experimental section. Furthermore, a preferred embodiment has the drug-containing microparticles imbedded in the (medicated) gel for local delivery.
  • CIJ confined impinging jets
  • F-127 E 99 -P 6 Q-E 99
  • F-87 E 61 -P 40 -E 6J
  • Both F127 and F87 consist of 70% polyethylene oxide) and 30% polypropylene oxide).
  • F-127 (ca.12,000 Da) gels at approximately 17 0 C for a 30% solution;
  • F-87 (ca. 7,700 Da) gels at around 40 0 C for a 30% solution.
  • the structures of the two gels Docket No. 788-74 (R7921) also differ; a BCC structure for F-87 and a FCC structure for F- 127.
  • the sol-gel transition temperatures can be altered between 17 0 C and 40 0 C by mixing the two polymers together at different weight ratios.
  • a series of solutions was prepared using different F87/F127 weight fractions at a fixed total copolymer concentration of 30% (w/v).
  • the sol-gel phase diagram of 30% (w/v) F87/F127 mixture solution at different F87 weight fractions in water is shown in Figure 2. It was observed that the lower critical separation temperature (LCST) of Pluronic mixture solutions increased with increasing F87 content at the fixed total copolymer concentration of 30% (w/v). It is noted that the 30% total concentration value is an arbitrary one.
  • the total polymer concentration can vary from just above the micelle overlap concentration to very high total polymer concentrations.
  • the sol-gel transition temperature could be controlled over a wide temperature range (from -17 0 C to -40 0 C) by varying the copolymer weight fractions. It should be noted that all sol-gel transitions are thermo-reversible.
  • the micelle formation induced by either a temperature change or a change in the concentration increment could be measured by an abrupt increase in the scattered intensity by means of laser light scattering.
  • the critical micelle temperature (cmt) defined as the temperature at which the light-scattering intensity departs significantly from the baseline intensity contributed only by unimers.
  • Figure 3a shows typical static (or time-averaged) light-scattering intensity results for the determination of the critical micelle temperature (cmt) of F87 solutions at various concentrations. This figure illustrates the concentration dependence of critical micelle temperature. It was seen that the polymer solution with higher concentrations had lower critical micelle temperatures. The dependence of critical micelle temperature on polymer solution concentration is plotted in Figure 3b.
  • micellar aggregation number N w CN A a 3 /(2M WI ), where
  • N A , and Mw t are polymer solution concentration (g/mL), Avogadro number and molecular weight of the polymer, respectively.
  • a 30% (w/v) F87 aqueous solution at 42 0 C has an aggregation number of about 57 which is a little bit smaller than that (-65) of 21.2% F127 solution in its gel-like state.
  • the ordered structures of F87/F127 mixed solution at different weight ratios in their gel-like states were also preliminarily investigated by the SAXS technique. Two kinds of ordered structures were observed when the F 127 content (weight percent) in the mixed solution was changed from 0 to 30%. At lower Fl 27 content (less than 15%), the mixed solution showed a body-centered cubic structure (BCC). When the Fl 27 content reached 20%, a transition state was observed. Further increasing the F 127 content in the mixed solution, face-centered cubic structure (FCC) was observed, instead of the BCC structure.
  • FCC face-centered cubic structure
  • Figure 5 shows the change of the lattice constant with the F87 content in the mixed solution.
  • the lattice constant could be tuned from 17 to 28 nm by changing the F87/F127 ratios in the solution at a fixed total concentration of 30% (w/v).
  • FIG. 6 Schematic representations of F87 and F 127 triblock copolymer micelles in the cubic structures are shown in Figure 6.
  • the dark blue region represents the hydrophobic core of poly(propylene oxide) in the micelle.
  • the green region represents the dimensions of the hydrated poly(ethylene oxide) chains.
  • the green gradient represents the decaying radial density profile of the poly(ethylene oxide) shell.
  • the bright white region is the interstitial region relatively free of overlap between micelles.
  • the solution viscosity is an important parameter for Pluronic applications.
  • a polymer solution with low viscosity will allow the rapid loading into devices under low applied pressure and hence can facilitate easier full automation of the system.
  • Figure 7 shows the temperature dependence of the zero-shear rate viscosity of 30% (w/v) F87/F127 mixed solution at a weight ratio of 1 :2 in aqueous solution. With increasing temperature from zero 0 C, the solution viscosity first decreased slightly at low temperatures and reached a minimum value of about 44 cP at 10 0 C. This is due to the shrinkage of coil size arising from an increase in the P block hydrophobicity. Further increasing the temperature will cause the viscosity to increase until the gel-like state is reached.
  • FIG. 8 illustrates the temperature dependence of the zero-shear rate viscosity of 30% (w/v) F87/F127 mixed solutions (weight ratio of 1:2) at different Docket No. 788-74 (R7921) concentrations of Iidocaine. With increasing weight concentration of lidocaine, the gelation temperature was shifted to a higher value. Understanding of such a shift under clinical conditions will be of importance.
  • a drug release test was conducted and the following experimental procedures were used. Solid lidocaine hydrochloride was mixed with 30.0 wt% Pluronics of different ratios. These samples were vortexed every hour, and refrigerated until complete dissolution. The drug release profile was tested by using a dialysis membrane (Spectra/Por ® membranes, MWCO: 3,500 produced by Spectrum Labs). The dialysis tube was immersed in a phosphate buffer solution at a pH of 7.2 and 37 0 C. The amount of lidocaine delivery, representing the fastest diffusion limit of lidocaine release, is shown in Figure 9, when compared with actual applications where the gel will not likely be in a large pool of body fluid.
  • the release rate is fairly linear and the amount released is essentially independent of the ratio of F87/F127, implying that the amount released is proportional only to the total copolymer concentration. While the variation of the F87/F127 ratio can change the sol-gel transition temperature, it is the total amount of polymers present that can affect the release rate as the release rate depends on the copolymer mesh size. Thus, for fast release, we have a pathway to make suitable adjustments, in both the release rate and the gel transition temperature.
  • lidocaine is not covalently bounded to the Pluronics copolymer, all of the soluble lidocaine will be released in time, in a linear release rate with the rate depending mainly on the diffusion coefficient of lidocaine in the polymer network and the concentration gradient of lidocaine between the gel and the surrounding medium.
  • the oil phase was prepared by dissolving 2.5 wt% of Span 80 surfactant in mineral oil, microparticles were obtained by dropping alginate/chitosan/lidocaine solution into the oil phase simultaneously.
  • Span 80 surfactant Span 80 surfactant
  • the in vitro release test was also carried out in a dialysis tube containing emulsified solution immersed in 100 mL phosphate buffer (pH 7.2) at 37 0 C.
  • lidocaine/Pluronic smart gels which can be used to administer the short term delivery (to 12 hours)
  • lidocaine PC microparticles which can be used to administer the long term delivery (to tens of days). It is logical to use the blends of two systems to control the sol-gel transition temperature as well as the release rate and release profile, as microparticles of different sizes can be incorporated into the lidocaine- containing pluronic solution prior to administration.
  • polymer blends of Pluronic block copolymers (EPE pluronics with different E and P contents as well as PEP pluronics) are used as a thermo-reversible sol-gel drug carrier.
  • EPE pluronics with different E and P contents as well as PEP pluronics are used as a thermo-reversible sol-gel drug carrier.
  • the incorporation of relatively more hydrophobic PEP component is important as the mixed gels can remain water-resistant for a long period of time, i.e., the medicated gels will not become soluble during the period of lidocaine controlled delivery.
  • Pluronic copolymers F127 (E 99 P 69 E 99 , with subscripts denoting the number of segments in the polymer chain and E, P representing oxyethylene and oxypropylene, respectively) and F87 (E 61 P 40 E 61 ) can be purchased directly from the BASF Corporation.
  • the first goal of this task is to find the solution behavior for optimal use in the body.
  • the general behavior of the pluronics can be represented schematically in Figure 11.
  • the sol-gel transition behavior depends on the nature of the solvent, the copolymer chain length and block length ratio, and the total copolymer concentration. Instead of synthesizing different copolymers, we simply will use blends of such copolymers.
  • Pluronic copolymers will be mixed in different concentrations and ratios in order to give a variety of gelation points.
  • the first set of solutions will be made of just one type of copolymer with Docket No. 788-74 (R7921) water.
  • the concentrations of the solutions will range from 10.0%-50.0% (w/v) with intervals of 5.0%. Therefore, there will be two solutions for each percentage; one for F 127, and one for F87.
  • the solutions will be tested in a water bath at 37 0 C to see whether or not the solutions can gel at this temperature.
  • the exact sol-gel transition temperatures of the solutions will be noted by changing the temperature of the water bath between 10 0 C and 7O 0 C.
  • the solutions will be characterized by using small-angle X-Ray scattering (SAXS) at the National Synchrotron Light Source (NSLS), Brookhaven National Laboratory (BNL) to determine whether the gel structure is body-centered cubic (BCC) or face-centered cubic (FCC), as shown in Figure 6.
  • SAXS small-angle X-Ray scattering
  • NLS National Synchrotron Light Source
  • BNL Brookhaven National Laboratory
  • Figure 12 shows the phase diagram at 37 0 C for copolymer blends of F127 and F87 in water.
  • the transition lines will change when the fluid properties change, such as an addition of salt or lidocaine in the solvent mixture ( Figure 8). However, the overall phase transition behavior should remain relatively constant.
  • thermo-reversible viscosity In order to determine the thermo-reversible viscosity and to make sure that the solution is a fluid at low viscosity levels over a temperature range which includes room temperature and becomes a gel at or near 37°C (body temperature) and that a gel formed when the droplets reach body temperature the polymer blends at different F127/E87 ratio and a fixed total polymer concentration are systematically measured.
  • Figure 8 shows a typical viscosity versus temperature curve for one of those pluronics blends without lidocaine.
  • the viscosity is suitable for spraying the solution into droplets.
  • the small droplets can convert into a sticking gel when the droplets touch the tissue surface at body temperature.
  • the next step is to prepare a polymer solution containing lidocaine.
  • Solid lidocaine hydrochloride is mixed with 30-50 (w/v)% Pluronics of Docket No. 788-74 (R7921) different ratios so that the final lidocaine hydrochloride concentration will be 2-5(w/v)%.
  • Pluronics of Docket No. 788-74 R7921
  • These samples go through the same established procedure, i.e., vortexed every hour, and refrigerated until complete dissolution.
  • the drug release profile is then tested by using a dialysis membrane (Spectra/Por® membranes, MWCO: 3,500 produced by Spectrum Labs).
  • the dialysis tube will be immersed in a phosphate buffer solution at a pH of 7.3 and 37 0 C. It is noted that in this in vitro experiment, the amount of lidocaine delivery, represents the fastest diffusion limit of lidocaine release and the sol-gel transition temperature should also depend on the amount of lidocaine used. The sol-gel transition temperature shift, however, is expected to be in the right direction, i.e., with release of lidocaine, the sol-gel transition temperature will be lowered, further stabilizing the gel.
  • the polymer network size (and thus the lidocaine diffusion rate from the polymer network) can be designed to provide a proper mesh size by varying the total Pluronics copolymer concentration, while the sol-gel transition temperature depends mainly on the copolymer block length and block ratio as well as the lidocaine content.
  • Lidocaine hydrochloride the water-soluble and anesthetically more active form
  • Pluronic copolymer blends of Fl 27 and F87 Pluronic copolymer blends of Fl 27 and F87. Its temperature-dependent viscosity behavior can be fine-tuned and is suitable for different forms of delivery, including the spray delivery format.
  • thermo-sensitive polymer network can be maintained over longer time periods. It is noted that a photo-sensitive or thermo-reactive cross-linking agent can be added to further stabilize the integrity of the gel, in the presence of excessive amount of fluid flow. Docket No. 788-74 (R7921)
  • At least two forms of copolymers can be used to fabricate at least two types of thermo-sensitive gels.
  • These two type of gels although differ in morphology, are formed from the same building blocks and therefore should be miscible, permitting fine-tuning of the delivery vehicle.
  • medicated PC particles described next
  • the basic polymer mesh size and the sol-gel transition temperature can be adjusted independently through the matrix components.
  • the present invention describes a simple but intricate and flexible time delay or time-release delivery system for different types of therapeutic agents.
  • lidocaine in either alginate/chitosan PC, or their mixtures using the recently developed process termed Flash NanoPrecipitation are produced.
  • the therapeutic agent used in the present invention is lidocaine, it is within the scope of the invention to use other therapeutic agents in place of or in addition to lidocaine.
  • This process utilizes a confined impinging jets (CIJ) mixer as shown Docket No. 788-74 (R7921) schematically in Figure 13.
  • CIJ mixer two high velocity jets (with speeds up to a few meters/sec) impinge each other through a coUinear aligned capillary (marked light green).
  • the separation length scale of the incoming fluids is rapidly reduced by these two equal momentum collinear jet streams.
  • the size of the mix chamber (marked dark green) is sufficiently large to allow for the formation of an impingement lane, but confined to avoid significant "bypassing", i.e., without encountering the other opposing stream.
  • the characteristic mixing time can be adjusted by changing the speed of the impinging jets.
  • the mixing time D T 1 , ⁇ can be as short as a few milliseconds. If the jet velocities in the region that have characteristics of a turbulent-like flow pattern within the cylindrical mixer, it will further reduce the mixing time.
  • the reaction products after the rapid mixing will be guided through a long tube runner (marked blue) to ensure that the two opposing steams were fully reacted before final collection.
  • one impinging stream contains cationic chitosan and slightly cationic lidocaine, while the other stream contains anionic alginate.
  • Both PC components i.e., chitosan and alginate
  • the weak formed complex and only slightly insoluble lidocaine-alginate complex can be entrapped within the particle.
  • the process is based on the supposition that lidocaine will react with alginate and form the complex that can be incorporated into the PC. If the mixing is faster than the induction time for precipitation, the process will be in a "homogeneous" condition, i.e., the effect of mixing is not convoluted with the precipitation times.
  • the size of the formed nanoparticles can be designed by carefully choosing the concentration of drug, which tunes the ⁇ ng , and the properties of PC which determine ⁇ agg .
  • both the particle size and the amount of drug inclusion can be controlled, without the use of surfactants.
  • the detailed balance of the two competing processes will be determined by drug concentration, kinetics of PC formation and lidocaine complex formation, but the formed particles will be expected to have a narrow size distribution.
  • the present invention takes advantage of the unique properties of lidocaine in the impinging jet approach, by making it less soluble either when forming a complex with alginate or when forming 'endless chains'. By using this technique we can avoid the use of Span 80, to form more uniform sized lidocaine-containing particles, and to have the pathway for easier scale-up productions.
  • Span 80 narrow size distribution micro- and nano-particles can be mixed together to tailor-design the drug release profile.
  • thermo-reversible smart gels using two rat models: the limb-movement test and the well-established tail-flick test can be carried out.
  • the primary objective of this task is to determine the relationship between the efficacy of medicated gel in a period of one week and the appropriate dosage of lidocaine loading.
  • the initial drug concentration of the gel, to be applied to the rat models will be formulated based on the release profile to simulate what is required for systemic treatment in human patients after surgery. As the local concentration of the drug around the injured site can be very high, we expect that the drug concentration can be significantly lowered in actual applications, but this will be determined by the in vivo tests for a specific target.
  • the present invention directed to smart gel formulations has the following advantages.
  • the materials are biocompatible, bioabsorbable or/and biodegradable.
  • the formulation is ideal for local pain-management on injured sites (internal and topical).
  • thermo- reversible gels can be designed over a very wide range from hours to days.
  • Pluronic block copolymer blends with thermally reversible gelation behavior near body temperature and water resistant characteristics during drug delivery, will be particularly useful for localized deployment (e.g. aerosol spraying, injection and topical coating)
  • the particle formation can be achieved without the use of surfactants.
  • the amount of drugs, together with particle size will provide the essential variables needed to tailor design both the amount and the long-term release rate of drugs.
  • the dorsum skin defect model can offer a sophisticated method to evaluate the efficacy of the pain-releasing device.
  • the application of strain gauges to the extremities of the rats to quantify movement shall enable us to gather more analytical data.
  • the Dorsum Skin Defect Model also known as "the hot plate test” can be used to test Sprague Dawley (SD) rats as an animal model of mammalian vertebrates that appear to experience pain in a similar fashion to humans. Initially, rats will be acclimated to sling restraints that position the rat so their paws are just able to touch the floor. This will be done by placing each rat into the sling for several minutes on multiple occasions until they become accustomed to the device as shown by minimal reaction or movement during the time they are in the sling. After acclimation, strain recorders will be attached to the rat legs and to the sling, with baseline movement levels established for each animal.
  • SD Sprague Dawley
  • the SD rats will then be divided into 3 groups: (1) a sham surgery control, (2) a group receiving surgery and no treatment, and (3) a group receiving surgery and medicated gel.
  • the initial dosage of Lidocaine in the trial membranes during in vivo study will be calculated based on Docket No. 788-74 (R7921) the clinical dosage (100 mg) given to a patient of 70 Kg, which can be viewed as the upper limit.
  • the optimized dosage will be determined through this in vivo study.
  • the surgery will consist of removal of 1 square centimeter piece of skin on the back, just behind the shoulder blades.
  • the rats will be allowed to recover from anesthesia and then placed in the sling and their level of analgesia assessed. This will be done by exposing the surgical site (or skin in the same location for sham controls) to an infrared light source that will heat the area and intensify any discomfort that might be present. The reaction of the rat to this stimulus will be recorded on videotape as well as via strain gauges attached to the rat legs and to the sling. The amount and force of movement as the animal attempts to escape the light source should correlate to the level of discomfort being experienced by the rat. The sling will prevent random locomotion of the animal that may confound interpretation of their escape response.
  • the rat tail-flick test can be adapted from previously established models, in which the level of discomfort experienced by the rats will be indirectly evaluated from physical movements.
  • Sprague-Dawley rats 200-250g are trained to remain stationary while being immobilized in a restraint unit (Lomir Biomedical, Inc.).
  • the analgesic groups are further divided by the composition of the controlled drug-delivery systems.
  • a predetermined distal portion of their tails are subjected to thermal stimulation by a focused beam of high-intensity light, and the response times to these stimuli are measured; the extent of movement are also recorded by the number of tail flicks, or any other bodily flinches, that occur.
  • the light source is applied for no longer than 30 seconds, or for 10 seconds after escape movements are noted.
  • the analgesic groups that will have the drug medication sprayed on will have the thermal stimulus applied at 30 minutes post-surgery, 45 minutes, 1 hour, 3 hours, 8 hours, 24 hours and Docket No. 788-74 (R7921) daily up to 7 days to determine the longevity of efficacy of the applied controlled drug delivery systems. Terminal blood samples from each animal will be analyzed for indicators of pain and distress.
  • the present invention is directed to (1) Medicated smart gels, containing at least one therapeutic agent, such as lidocaine, (an anesthetic (model) drug in the present discussion) and mixtures of Pluronic-based block copolymers (poly(ethylene oxide)- ⁇ -poly(propylene oxide)-Z>- poly(ethylene oxide)) in water or buffer solution which can be fine-tuned by copolymer composition and concentration to possess thermo-reversible gelation behavior near the body temperature.
  • Such characteristics shall enable the sustained and controlled release of lidocaine molecules and can facilitate site deployment pathways (e.g. low pressure spraying, injection or topical coating).
  • microscopic bioabsorbable medicated particles micron and submicron sizes
  • polyelectrolyte complex formation e.g., mixtures based on chitosan and alginate
  • the combination of these two technologies offers a powerful means to prepare effective pain-management media for site-specific treatments as well as treating other diseases which require a steady administration of drugs with a beneficial delay-release.

Abstract

A thermo-reversible gel containing sub-micron size particles containing at least one therapeutic molecule and alginate polyelectrolyte-complex, chitosan polyelectrolyte-complex or mixtures thereof is described herein. One particular therapeutic molecule that can be used in the thermo-reversible gel is lidocaine.

Description

NANOSTRUCTURED SMART GEL FOR TIME RELEASE DRUG DELIVERY
PRIORITY
This application claims priority to provisional application serial number 60/844,575, filed with the U.S. Patent and Trademark Office September 14, 2006, the contents of which are incorporated herein by reference.
BACKGROUND
1. Field of the Invention
The present invention is directed to a unique class of nanostructured smart gels for site- specific pain management.
2. Background of the Invention
Chronic pain has been traditionally defined as pain persisting for at least 3-6 months but can now be defined as pain that extends beyond the period of tissue healing and/or with low levels of identified pathology that are insufficient to explain the presence and/or extent of the pain. The persistence of chronic pain can affect many different aspects of personal life, ranging from the physiological to societal and financial changes. Chronic pain afflicts 24% of all Americans, most prominently affecting women who are over 50 years of age. Chronic pain patients typically average 12.9 trips to a General Practitioner per year - as opposed to the average 4.2 trips per year for the general community. Overall, the economic burden of the lost productivity due to chronic pain totals approximately $86.2 billion per year for the United States, a figure sure to exponentially rise as the "Baby Boomer" generation ages and no definitive therapeutic treatments have been discovered.
Lidocaine is one of the most widely used anesthetic drugs today, especially during surgery and dental procedures. Lidocaine was developed by the Swedish scientist Nils Lδfgren in 1943. An additional use of lidocaine is as an anti-arrhythmic agent. The drug works by inhibiting the stimulants needed to initiate neuronal impulses to the brain, resulting in the loss of pain. As a Docket No. 788-74 (R7921) topical drug, lidocaine has a relatively short half-life of only 1.5-2 hours in an intravenous injection because it is quickly metabolized by the liver (due to the presence of an amide group). Even though the time frame in which lidocaine works is extremely short, it is commonly used as the local anesthetic of choice among professionals due to its hypoallergenic quality. Allergic reactions to lidocaine are extremely rare, and if they do occur, it is usually due to one of the preservatives found in the dose vials. This present invention uses a thermally responsive carrier for controlled delivery of a known and desired amount of lidocaine to an injured site with pre-designed release profile over a flexible time period from many hours to 10-15 days. The thermo-reversible sol-gel carrier, which is a liquid at room temperature and a gel at body temperature, allows versatile deployment methods to be implemented during medical procedures.
SUMMARY OF THE INVENTION
An aspect of the present invention is to substantially solve at least the above problems and/or disadvantages and to provide at least the advantages below. The present invention provides a way to provide the pain relief attributes of lidocaine as well as other therapeutic agents, including as prilocaine, bupivacaine, ropivacaine molecules, analogs and mixtures thereof, to a patient for a sustained period of time without affecting the attributes of the drug.
In one embodiment, the invention utilizes smart gels as a drug delivery system. The therapeutic agent(s) can be added directly to the smart gel itself for faster release, as well as to particles incorporated into the smart gel for sustained release. The (sub-)micron size particles of the invention contain at least one therapeutic agent, such as lidocaine, prilocaine, bupivacaine, ropivacaine molecules, their analogs or their mixtures. The particles are based on alginate/chitosan polyelectrolyte-complex (PC), chitosan polyelectrolyte-complex (PC) or mixtures. The polyelectrolyte complex.alginate/chitosan forms one ion pair; but there are many others as well that Docket No. 788-74 (R7921) can be utilized in accordance with the invention so long as they are oppositely charged and nontoxic, preferably with components already approved by the FDA. The particle size ranges from about 10 am to about 1 micon; the alginate/chitosan ratio can be from 100/0 to 0/100. The present invention provides a time-release system that allows the desired prolonged drug delivery in pain management.
In one embodiment, he smart gel of the present invention is a a thermo-reversible gel composition. Said gel can contain therapeutic agent(s) added directly thereto for faster release, as well as to particles incorporated into the smart gel for sustained release. When (sub-) micron size particles are employed, said particles contain therapeutic agents, such as lidocaine prilocaine, bupivacaine, ropivacaine molecules, their analogs or their mixtures; and alginate polyelectrolyte- complex (PC), chitosan polyelectrolyte-complex (PC), and mixtures thereof. The thermo- reversibility is obtained by changing the temperature of the solution and breaking the hydrogen bonds that are formed at a lower temperature and thereby causing the sol to gel and release the therapeutic agent is slowed down by the polymer network, whose mesh size can partially be controlled by the copolymer concentration. The diffusion of therapeutic agent through the polymer network is much slower when compared with that in a polymer solution and thereby in slowing down the time release of the therapeutic agent.
Another embodiment of the present invention is directed to a smart gel including a thermo- reversible Pluronic hydrogel having a series of (E) and (P) triblock copolymers with a general formula of ExPyEx, wherein E is a hydrophilic molecule and P is a molecule that is capable of temperature-dependent hydrogen bonding with water (near the body temperature) and therefore is hydrophilic at low temperatures, but becomes hydrophobic at higher temperatures due to the break down of hydrogen bonds between the P blocks and water. The value of x can range from about 1 to about 300 and the value of y can range from about 1 to about 300. Docket No. 788-74 (R7921)
Still another embodiment of the present invention is directed to a smart gel comprising a thermo-reversible Pluronic hydrogel comprising a series of (P) and (E) triblock copolymers with the general formula Of PxEyPx. As with the embodiment described above the value of x can range from about 1 to about 300 and the value of y can range from about 1 to about 300. The switching of roles between E and P, i.e., from ExPyEx to PxEyPx, changes the morphology of the polymer network formed.
In a preferred embodiment of the present invention the thermo-reversible Pluronic hydrogel compositions described above can be blended together to produce a single composition having combined properties. The thermo-reversible Pluronic hydrogel blend comprising a series of (E) and (P) triblock copolymers with the general formula of ExPyEx, wherein E is a hydrophilic molecule and P is a molecule that is capable of a stronger temperature dependent hydrogen bonding with water and therefore is hydrophilic at low temperatures, but becomes hydrophobic at higher temperatures due to the break down of hydrogen bonds between the P blocks and water. The value of x can range from about 1 to about 300 and the value of y can range from about 1 to about 300. The composition also includes a series of (P) and (E) triblock copolymers with the general formula of PxEyPx, wherein E is a hydrophilic molecule and P is a molecule that is capable of a stronger temperature dependent hydrogen bonding with water (near body temperature) and therefore is hydrophilic at low temperatures, but becomes hydrophobic at higher temperatures due to the break down of hydrogen bonds between the P blocks and water having a value of x that ranges from about 1 to about 300 and a value of y that ranges from about 1 to about 300. The composition contains at least one therapeutic agent, such as lidocaine, and results in a time-released composition. Docket No. 788-74 (R7921) BRIEF DESCRIPTION OF THE FIGURES
The above and other objects, features and advantages of the present invention will become more apparent from the following detailed description when taken in conjunction with the accompanying drawing in which:
Figure 1 is the structure of lidocaine hydrochloride;
Figure 2 is the Sol-gel phase diagram of aqueous 30% (w/v) F87/F127 mixture solutions;
Figure 3a shows typical static light scattering results for the determination of cmt of F87 solutions at various concentrations;
Figure 3b graphically shows the dependence of critical micelle temperature on polymer solution concentration;
Figure 4 illustrates a typical SAXS intensity profile of 30% (w/v) F87 aqueous solution at 420C;
Figure 5 shows the change of the lattice constant with the F87 content in the mixed solution;
Figure 6 schematically shows representations of F87 and F127 triblock copolymer micelles in the cubic structures;
Figure 7 graphically illustrates the temperature dependence of the zero-shear rate viscosity of 30% (w/v) F87/F127 mixed solution at a weight ratio of 1 :2 in aqueous solution;
Figure 8 graphically illustrates the temperature dependence of the zero-shear rate viscosity of 30% (w/v) F87/F127 mixed solutions (weight ratio of 1:2) at different concentrations of lidocaine;
Figure 9 graphically illustrates the amount of lidocaine delivery, representing the fastest diffusion limit of lidocaine release;
Figure 10 graphically illustrates the typical release profiles from the solution containing 2 wt% chitosan (10 mL), 4.4 wt% alginate (5 mL) and mineral oil (10 mL); Docket No. 788-74 (R7921)
Figure 11 schematically represents the general behavior of the pluronics of the present invention;
Figure 12 shows the phase diagram at 37 0C for copolymer blends of F127 and F87 in water; and
Figure 13 schematically shows a confined impinging jets (CU) mixer.
DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS
Preferred embodiments of the present invention are described herein with reference to the accompanying drawings. The therapeutic value of many different therapeutic agents or drugs has only lasted for short periods of time and requires frequent administration of the drug to maintain a desired condition. This is surely the case with pain relief medication such as lidocaine.
There are two forms of lidocaine (2-diethylamino-N-(2,6-dimethylphenyl) acetamide): lidocaine hydrochloride (lidosalt), as shown in figure 1 and lidocaine base (lidobase). Lidocaine hydrochloride with the formula Of Ci4H22N2O • HCl is the anesthetically active form and is soluble in water, whereas lidobase with the formula Of Ci4H22N2O is not soluble in water and anesthetically much less active. Lidocaine hydrochloride is predominately marketed by AstraZeneca as Xylocaine®.
Lidocaine hydrochloride has a very interesting structure in water. The hydrate microcrystal theory of anesthesia, advanced by Linus Pauling is closely related to the "iceberg" theory of ionic solutions and hydration of proteins. The ordered arrangement of water molecules around solute ions and protein side chains is considered as part of the clathrate structure. Later studies have shown that local anesthetics form a hydrogen bonded complex with a receptor in the membrane. Lidocaine, a sodium ion channel blocker or nerve block, is a local anesthetic. A lidocaine cation can donate two protons and accept one. The crystal structure of lidocaine indicates that adjacent chains are held Docket No. 788-74 (R7921) together by chloride ions, each of which accepts an aqueous proton from one chain and an amino proton from the other. The double chains are held together by van der Waals forces. The structure is fully hydrogen bonded and 'endless chains ' of lidocaine cations are produced by water molecules.
Infrared spectra of lidocaine cation with different anions indicate strong interactions between nitrogen and hydrogen and a small anion associated with an intense force field such as chloride. On the other hand, with a large polyatomic anion associated with a weaker peripheral force field such as hexafluoroarsenate, the N+-H stretching frequency is higher than that with chloride ion. The neutral form of lidocaine is about five times less effective when compared with the cationic form. With the physiologically active form of lidocaine being the cation, the choice of an appropriate anion may also influence the rate of drug delivery from the gel. The spectrophotometric results reported herein at 263 ran indicate significant deviations from Beer's law above 0.01 M solutions of lidocaine hydrochloride, suggesting strong π - π interactions at higher concentrations. It also suggests the possibility of 'endless chain' formation for lidocaine in solution as well as in the hydrated crystal formation.
In a recent paper on sustained release of lidocaine from Poloxamer (an alternative name for Pluronics) 407 gels, it was assumed that lidocaine, because of its charge, is located in the outer aqueous region of the gel rather than in the hydrophobic interior of the micelles. This supposition has not been verified experimentally. Moreover, crystal data suggest that it is possible for lidocaine cations to enter the micelle by hiding its charge with the formation of 'endless' chains in the presence of chloride ions and hydrogen bonding with water. This observation necessitates a further need for data from laser light scattering as well as small angle and wide angle X-ray scattering.
The electrochemical impedance measurements of lidocaine hydrochloride reported herein indicate negative differential resistance at some potential, suggesting possible tunneling of electrons Docket No. 788-74 (R7921) through the self assembled pathway or through the ordered arrangement of water molecules in between the chains near the electrode double layer.
Nanostructures due to Self- Assembly of Pluronic Molecules
Pluronic hydrogels are often used in pharmaceutical and biotechnological industries because of its low toxicity, little to no immune response and unique temperature-dependent viscosity. Hydrogels are the closest synthetic materials to emulate living tissue due to its ability to imbibe large amounts of water as well as its soft consistency. The amount of water that a hydrogel can take in is limited by the force of mixing as well as the retractive force of the polymer. Therefore, there is an upper limit as to how much of the hydrogels will, in fact, be dissolved. These factors greatly influence its low immune response.
Pluronic hydrogels are a series of ethylene oxide (E) and propylene oxide (P) triblock copolymers with the general formula of ExPyEx, which can be considered as amphiphilic non-ionic surfactants. At higher enough concentrations, their aqueous solutions exhibit liquid-like properties at low temperatures and gel-like properties at higher temperatures because the E blocks are consistently hydrophilic, which allows for the copolymer to be dissolved in water. The middle P block, on the other hand, is hydrophilic at low temperatures, but hydrophobic at higher temperatures due to the break down of hydrogen bonds between the P blocks and water. This distinctive property of thermally responsive Pluronic hydrogels makes it easy to administer to patients. The solution can be given through injection at a low temperature, and once the liquid enters the body, it would immediately gel. The physically cross-linked gel would stay essentially in place right at the zone of treatment without having to worry about the drug impacting other areas of the body.
The thermally responsive gelation behavior of Pluronic copolymers changes with the E/P ratio and the total chain length. There are two main types of structures that these polymers can form Docket No. 788-74 (R7921) when the solution gels. One is the body centered cube (or BCC), and the other is the face centered cube (FCC) phase. The exact function that these structures play in drug delivery could use further study.
Biodegradable Polyelectrolvte Complexes
Developments on the formation of oppositely charged polyelectrolvte complexes are slightly different from polyelectrolyte-surfactant complexes (PSCs) that have been extensively described in the literature. PSCs are microphase-separated systems containing hydrophilic domains (with surfactant head groups and polyelectrolvte charges) and hydrophobic regions (with surfactant tails and hydrophobic parts of the polyelectrolvte). The two most important driving forces for the self- assembly of surfactant molecules in PSCs are: (1) electrostatic interactions between the charged components and (2) hydrophobic interactions between the polymer backbone and the alkyl chains of the surfactant. These well-defined supramolecular structures of the polyelectrolyte-surfactant complexes have unusual mechanical, optical, electrical and biological properties, which could be very useful in the medical (e.g., gene therapy), cosmetic, food, painting, coating* enhanced oil recovery, and other technologies.
In polyelectrolvte complexes (PCs), the oppositely charged polyelectrolytes are neutralized via electrostatic interactions, forming ion pairs and rendering hydrophobic regions, such as in the neutralization of negatively charged alginate and positively charged chitosan, in which both components are degradable in water (alginates are produced by brown seaweeds - phaeophyceae, mainly laminaria; and chitosan can come from a bacterial polysaccharide). In such cases, the drug molecules can be entrapped in the complexes for a prolonged period of time without diffusing out of the complex domain, as has been demonstrated by our preliminary experiments. Docket No. 788-74 (R7921)
Microparticle Formation Techniques
Biodegradable microparticles have been routinely used in oral delivery systems and, even more often, in subcutaneously injected delivery systems because they can be administered to a variety of locations in vivo through a syringe needle. The release rate of drugs from biodegradable microparticles can be controlled by a number of factors, such as polymer biodegradation kinetics, physicochemical properties of polymers and of drugs, the size and size distribution of microparticles. For typical small molecule therapeutics, drug release often exhibits an initial 'burst' phase during which a significant fraction (5-50%) of the encapsulated compound is released in a short time (<8 h). The burst is often undesirable because the initial dosage may result in toxicity (due to over dosage within a very short time period) or other side-effects. Among all parameters affecting the drug release rate, microparticle size is an important determinant of drug release rate. Larger spheres generally release encapsulated compounds more slowly and over longer time periods. Thus, controlling microparticle size provides an effective and relatively simple pathway for control of drug release. Numerous studies have been conducted to determine the effects of microparticle size on drug release. Fabrication of biodegradable polymer microparticles with precise size and size distribution control provides a means for enhanced control of drug delivery rates.
The ability to control delivery kinetics is important for many applications. For example, frequent administrations needed for highly potent drugs with narrow therapeutic windows could be replaced by using a system capable of delivering the drug at a constant or predetermined rate for a prolonged time period. A variety of drugs, regardless of their molar mass and water solubility, can be loaded into biodegradable microparticles using different preparation techniques. However, polydispersed microparticles generated by conventional fabrication techniques must often be filtered to obtain the desired size range, causing a waste of polymer and preloaded drugs. Docket No. 788-74 (R7921)
Furthermore, microparticles with a relatively small size normally exhibit poor encapsulation efficiency and result in an undesirably rapid release of loaded drugs.
There are several techniques available for the production of microparticles containing the desired drugs, such as the emulsion-solvent evaporation/extraction method, spray drying and phase separation. Each technique has its own advantages and limitations. The choice of a particular technique depends on the attributes of the polymer and the drug, the site of drug action and the duration of the desired therapy. For example, the emulsion-solvent evaporation/extraction methods have been demonstrated to encapsulate both hydrophilic and hydrophobic drugs. However, the resulting particles often contain surfactant(s) that cannot be removed.
Recently, several methods for preparation of monodispersed polymer microparticles with the aid of microfluidics and acoustic excitation have been developed. In a preferred embodiment, a microparticle system consisting of tailor-designed size distribution that can cover a broad range of drug release rates is more effective than a monodispersed microparticle system that can provide only a narrow range of drug release profile. Accordingly, a newly developed fabrication process was used to overcome the above concerns. The technique is termed Flash NanoPrecipitation using a confined impinging jets (CIJ) mixer, which is discussed in the experimental section. Furthermore, a preferred embodiment has the drug-containing microparticles imbedded in the (medicated) gel for local delivery.
Relations of Structure and Property in Pluronic Gels
Two types of hydrogels, described herein as examples, were used in our preliminary study, F-127 (E99-P6Q-E99) and F-87 (E61 -P40-E6J ). Both F127 and F87 consist of 70% polyethylene oxide) and 30% polypropylene oxide). F-127 (ca.12,000 Da) gels at approximately 17 0C for a 30% solution; F-87 (ca. 7,700 Da) gels at around 40 0C for a 30% solution. The structures of the two gels Docket No. 788-74 (R7921) also differ; a BCC structure for F-87 and a FCC structure for F- 127. The sol-gel transition temperatures can be altered between 17 0C and 40 0C by mixing the two polymers together at different weight ratios. A series of solutions was prepared using different F87/F127 weight fractions at a fixed total copolymer concentration of 30% (w/v). The sol-gel phase diagram of 30% (w/v) F87/F127 mixture solution at different F87 weight fractions in water is shown in Figure 2. It was observed that the lower critical separation temperature (LCST) of Pluronic mixture solutions increased with increasing F87 content at the fixed total copolymer concentration of 30% (w/v). It is noted that the 30% total concentration value is an arbitrary one. -Depending on specifications as to polymer network mesh size and mechanical properties required, the total polymer concentration can vary from just above the micelle overlap concentration to very high total polymer concentrations. The sol-gel transition temperature could be controlled over a wide temperature range (from -17 0C to -40 0C) by varying the copolymer weight fractions. It should be noted that all sol-gel transitions are thermo-reversible.
In dilute Pluronic solutions, the micelle formation induced by either a temperature change or a change in the concentration increment could be measured by an abrupt increase in the scattered intensity by means of laser light scattering. The critical micelle temperature (cmt), defined as the temperature at which the light-scattering intensity departs significantly from the baseline intensity contributed only by unimers. Figure 3a shows typical static (or time-averaged) light-scattering intensity results for the determination of the critical micelle temperature (cmt) of F87 solutions at various concentrations. This figure illustrates the concentration dependence of critical micelle temperature. It was seen that the polymer solution with higher concentrations had lower critical micelle temperatures. The dependence of critical micelle temperature on polymer solution concentration is plotted in Figure 3b. Due to the small hydrophobic part of F87, its critical micelle concentration (cmc) is relatively high at 25 0C. Relative excess scattered intensity and critical Docket No. 788-74 (R7921) micelle temperature is shown in Figure 3 (a) and the dependence of critical micelle temperature on polymer solution concentration of F87 aqueous solution is shown in Figure 3(b).
At high concentrations and high temperatures, the pluronic micelles in aqueous solution are densely packed, leading to the formation of a gel-like ordered structure. Figure 4 shows the typical small-angle X-ray scattering (SAXS) intensity profile of 30% (w/v) F87 in aqueous solution at 42 0C. Four peaks were observed with their peak positions following a simple mathematical relation of
1 : Λ/2 : -JΪ :2, which are the typical Bragg scattering pattern for the body-centered cubic (BCC)
structure. From the position of the first scattering peak sno, the cubic lattice constant a =
Figure imgf000014_0001
*J2 was calculated to be about 16.9 nm. The center-to-center distance of the two closest neighbor micelles is -Jϊa/2 = 14.6 nm. From the lattice constant of the BCC structure, the micellar aggregation number Nw could be further calculated out by the equation Nw = CN Aa3 /(2MWI), where
C, NA, and Mwt are polymer solution concentration (g/mL), Avogadro number and molecular weight of the polymer, respectively. A 30% (w/v) F87 aqueous solution at 42 0C has an aggregation number of about 57 which is a little bit smaller than that (-65) of 21.2% F127 solution in its gel-like state.
The ordered structures of F87/F127 mixed solution at different weight ratios in their gel-like states were also preliminarily investigated by the SAXS technique. Two kinds of ordered structures were observed when the F 127 content (weight percent) in the mixed solution was changed from 0 to 30%. At lower Fl 27 content (less than 15%), the mixed solution showed a body-centered cubic structure (BCC). When the Fl 27 content reached 20%, a transition state was observed. Further increasing the F 127 content in the mixed solution, face-centered cubic structure (FCC) was observed, instead of the BCC structure. With the introduction of F87 in F127, we not only can alter the gel transition temperature but also can change the lattice constant of the ordered structure. Docket No. 788-74 (R7921)
Figure 5 shows the change of the lattice constant with the F87 content in the mixed solution. The lattice constant could be tuned from 17 to 28 nm by changing the F87/F127 ratios in the solution at a fixed total concentration of 30% (w/v).
Schematic representations of F87 and F 127 triblock copolymer micelles in the cubic structures are shown in Figure 6. The dark blue region represents the hydrophobic core of poly(propylene oxide) in the micelle. The green region represents the dimensions of the hydrated poly(ethylene oxide) chains. The green gradient represents the decaying radial density profile of the poly(ethylene oxide) shell. The bright white region is the interstitial region relatively free of overlap between micelles. There are at least four qualitatively different domains, namely the condensed P micellar cores, the hydrated E micellar shells, the entangled E chains in the overlapping micellar shells between the closest micelles, and the water-rich interstitial gaps between micelles.
Viscosity Changes in Medicated Pluronic Gels
The solution viscosity is an important parameter for Pluronic applications. A polymer solution with low viscosity will allow the rapid loading into devices under low applied pressure and hence can facilitate easier full automation of the system. Figure 7 shows the temperature dependence of the zero-shear rate viscosity of 30% (w/v) F87/F127 mixed solution at a weight ratio of 1 :2 in aqueous solution. With increasing temperature from zero 0C, the solution viscosity first decreased slightly at low temperatures and reached a minimum value of about 44 cP at 10 0C. This is due to the shrinkage of coil size arising from an increase in the P block hydrophobicity. Further increasing the temperature will cause the viscosity to increase until the gel-like state is reached.
We have noted that the addition of lidocaine exhibited a distinct shift in the gelation temperature of Pluronic solutions. Figure 8 illustrates the temperature dependence of the zero-shear rate viscosity of 30% (w/v) F87/F127 mixed solutions (weight ratio of 1:2) at different Docket No. 788-74 (R7921) concentrations of Iidocaine. With increasing weight concentration of lidocaine, the gelation temperature was shifted to a higher value. Understanding of such a shift under clinical conditions will be of importance.
Drug Release Study in Medicated Pluronic Gels
A drug release test was conducted and the following experimental procedures were used. Solid lidocaine hydrochloride was mixed with 30.0 wt% Pluronics of different ratios. These samples were vortexed every hour, and refrigerated until complete dissolution. The drug release profile was tested by using a dialysis membrane (Spectra/Por® membranes, MWCO: 3,500 produced by Spectrum Labs). The dialysis tube was immersed in a phosphate buffer solution at a pH of 7.2 and 37 0C. The amount of lidocaine delivery, representing the fastest diffusion limit of lidocaine release, is shown in Figure 9, when compared with actual applications where the gel will not likely be in a large pool of body fluid. It should be noted that there are two advantages to this result, i.e., the release rate is fairly linear and the amount released is essentially independent of the ratio of F87/F127, implying that the amount released is proportional only to the total copolymer concentration. While the variation of the F87/F127 ratio can change the sol-gel transition temperature, it is the total amount of polymers present that can affect the release rate as the release rate depends on the copolymer mesh size. Thus, for fast release, we have a pathway to make suitable adjustments, in both the release rate and the gel transition temperature. Finally, as lidocaine is not covalently bounded to the Pluronics copolymer, all of the soluble lidocaine will be released in time, in a linear release rate with the rate depending mainly on the diffusion coefficient of lidocaine in the polymer network and the concentration gradient of lidocaine between the gel and the surrounding medium. Docket No. 788-74 (R7921)
Sustained Drug Release in Medicated Polyelectrolyte Complex Microparticles
In a separate study, we have prepared biodegradable polyelectrolyte-complex (PC) microparticles using oppositely charged alginate and chitosan biomacromolecules. The experimental procedures were as follows. Sodium Alginate (Mv = 60k - 106k Da) was dissolved in distilled water, and chitosan (deacetation degree > 85%; Mv.= 817k Da) was dissolved in 2 wt% acetic acid aqueous solution. Both were then mixed to form a 2 wt% polymer solution, where lidocaine hydrochloride was subsequently added. The water-in-oil emulsification method was used to prepare medicated microparticles. The oil phase was prepared by dissolving 2.5 wt% of Span 80 surfactant in mineral oil, microparticles were obtained by dropping alginate/chitosan/lidocaine solution into the oil phase simultaneously. We recognize that the surfactant molecules were also encapsulated in the microparticles by this method. This problem will be overcome by the use of Flash NanoPrecipitation method, which will be discussed later. The in vitro release test was also carried out in a dialysis tube containing emulsified solution immersed in 100 mL phosphate buffer (pH 7.2) at 37 0C. The typical release profiles from the solution containing 2 wt% chitosan (10 mL), 4.4 wt% alginate (5 mL) and mineral oil (10 mL) are shown in Figure 10. Results indicate that by using lidocaine-containing PC microparticles, a fairly linear release rate of up to 10 days can be achieved. It is conceivable that a longer term (e.g. a month) drug delivery profile can be obtained if the composition of the sample is optimized.
The above results demonstrate at least two flexible platforms: (1) lidocaine/Pluronic smart gels, which can be used to administer the short term delivery (to 12 hours) and (2) lidocaine PC microparticles, which can be used to administer the long term delivery (to tens of days). It is logical to use the blends of two systems to control the sol-gel transition temperature as well as the release rate and release profile, as microparticles of different sizes can be incorporated into the lidocaine- containing pluronic solution prior to administration. Docket No. 788-74 (R7921) In Vitro Study of Lidocaine Containing Thermo-reversible Smart Gels
In a preferred embodiment, polymer blends of Pluronic block copolymers (EPE pluronics with different E and P contents as well as PEP pluronics) are used as a thermo-reversible sol-gel drug carrier. The incorporation of relatively more hydrophobic PEP component is important as the mixed gels can remain water-resistant for a long period of time, i.e., the medicated gels will not become soluble during the period of lidocaine controlled delivery.
The specific application of such gels for lidocaine delivery has never been attempted. The goals of this task thus are two: (1) tailoring the sol-gel transition temperature at around the body temperature of human (35-4O0C), and (2) initial control of the lidocaine release rate up to several days. We have chosen the Pluronics copolymers for reasons already discussed; the experiments described below shall describe the essential features that can satisfy the expected method of delivery, the sol-gel transition temperature, the integrity of the polymer network in the presence of body fluids and the relatively short drug release profile.
Pluronic copolymers F127 (E99P69E99, with subscripts denoting the number of segments in the polymer chain and E, P representing oxyethylene and oxypropylene, respectively) and F87 (E61P40E61) can be purchased directly from the BASF Corporation. The first goal of this task is to find the solution behavior for optimal use in the body. In order to discover the desired sol-gel transition temperature, we note that the general behavior of the pluronics can be represented schematically in Figure 11. The sol-gel transition behavior depends on the nature of the solvent, the copolymer chain length and block length ratio, and the total copolymer concentration. Instead of synthesizing different copolymers, we simply will use blends of such copolymers.
Pluronic copolymers will be mixed in different concentrations and ratios in order to give a variety of gelation points. The first set of solutions will be made of just one type of copolymer with Docket No. 788-74 (R7921) water. The concentrations of the solutions will range from 10.0%-50.0% (w/v) with intervals of 5.0%. Therefore, there will be two solutions for each percentage; one for F 127, and one for F87. The solutions will be tested in a water bath at 37 0C to see whether or not the solutions can gel at this temperature.
Next, the exact sol-gel transition temperatures of the solutions will be noted by changing the temperature of the water bath between 100C and 7O0C. The solutions will be characterized by using small-angle X-Ray scattering (SAXS) at the National Synchrotron Light Source (NSLS), Brookhaven National Laboratory (BNL) to determine whether the gel structure is body-centered cubic (BCC) or face-centered cubic (FCC), as shown in Figure 6. Figure 12 shows the phase diagram at 37 0C for copolymer blends of F127 and F87 in water. It should be noted that the transition lines will change when the fluid properties change, such as an addition of salt or lidocaine in the solvent mixture (Figure 8). However, the overall phase transition behavior should remain relatively constant.
In order to determine the thermo-reversible viscosity and to make sure that the solution is a fluid at low viscosity levels over a temperature range which includes room temperature and becomes a gel at or near 37°C (body temperature) and that a gel formed when the droplets reach body temperature the polymer blends at different F127/E87 ratio and a fixed total polymer concentration are systematically measured. Figure 8 shows a typical viscosity versus temperature curve for one of those pluronics blends without lidocaine. In the fluid range, the viscosity is suitable for spraying the solution into droplets. The small droplets can convert into a sticking gel when the droplets touch the tissue surface at body temperature.
Having selected a suitable pluronics solution, the next step is to prepare a polymer solution containing lidocaine. Solid lidocaine hydrochloride is mixed with 30-50 (w/v)% Pluronics of Docket No. 788-74 (R7921) different ratios so that the final lidocaine hydrochloride concentration will be 2-5(w/v)%. These samples go through the same established procedure, i.e., vortexed every hour, and refrigerated until complete dissolution. The drug release profile is then tested by using a dialysis membrane (Spectra/Por® membranes, MWCO: 3,500 produced by Spectrum Labs). The dialysis tube will be immersed in a phosphate buffer solution at a pH of 7.3 and 37 0C. It is noted that in this in vitro experiment, the amount of lidocaine delivery, represents the fastest diffusion limit of lidocaine release and the sol-gel transition temperature should also depend on the amount of lidocaine used. The sol-gel transition temperature shift, however, is expected to be in the right direction, i.e., with release of lidocaine, the sol-gel transition temperature will be lowered, further stabilizing the gel.
The lidocaine release results, without the optimized F87/F127 composition in the presence of lidocaine, are shown in Figure 9. It is appropriate to again emphasize the two advantages to this medicated system, i.e., the release rate is fairly linear and the amount released is essentially independent of the ratio of F87/F127, implying that the amount released is proportional only to the total polymer concentration. While the variation of the F87/F127 ratio can change the sol-gel transition temperature, it is the total amount of polymers present that can affect the lidocaine release rate as the release rate depends mainly on the copolymer mesh size. Thus, for relatively fast release (i.e., up to days), we have a pathway to make suitable adjustments, depending on specific applications. To summarize, the use of Pluronics triblock copolymers can offer the following advantages.
• By blending copolymers of different block length and block ratio, we can design a sol-gel transition temperature so that the solution has a low viscosity at room temperatures for convenient delivery in a spray format, while at body temperature (-37 0C), the solution containing lidocaine becomes a gel. Docket No. 788-74 (R7921)
• The polymer network size (and thus the lidocaine diffusion rate from the polymer network) can be designed to provide a proper mesh size by varying the total Pluronics copolymer concentration, while the sol-gel transition temperature depends mainly on the copolymer block length and block ratio as well as the lidocaine content.
• Lidocaine hydrochloride, the water-soluble and anesthetically more active form, can be mixed directly with Pluronic copolymer blends of Fl 27 and F87. Its temperature-dependent viscosity behavior can be fine-tuned and is suitable for different forms of delivery, including the spray delivery format.
• Finally, we note that there are two forms of micelles that can be formed in Figure 11, depending on the copolymer chain sequence (EPE or PEP). We have so far discussed the advantages of only the EPE-type tri-block copolymers. By examining the self-assembly behavior of the PEP- type tri-block copolymers, flower-like micelles are formed. With the middle soluble block being longer, dangling E-chains with open-ended P-blocks should appear. The flower-like micelles have the dangling P-blocks that can act as physical cross-linking points and form an open network. Thus, we can take advantage of this self-assembly behavior to form a physical gel that has a different property from the gel-like structure formed mainly by the overlap of E-chains in the shell of core-shell micelles, as also shown in Figure 11. The open network has physical cross-linking points and cannot be dissolved easily even in an open system with large amounts of soluble solvent for the E-chains. Thus, the integrity of the thermo-sensitive polymer network can be maintained over longer time periods. It is noted that a photo-sensitive or thermo-reactive cross-linking agent can be added to further stabilize the integrity of the gel, in the presence of excessive amount of fluid flow. Docket No. 788-74 (R7921)
In summary, at least two forms of copolymers can be used to fabricate at least two types of thermo-sensitive gels. These two type of gels, although differ in morphology, are formed from the same building blocks and therefore should be miscible, permitting fine-tuning of the delivery vehicle. Of course, with the incorporation of medicated PC particles (described next), the gelation and rheological properties of the polymer solution would change slightly. The basic polymer mesh size and the sol-gel transition temperature can be adjusted independently through the matrix components. Thus, the present invention describes a simple but intricate and flexible time delay or time-release delivery system for different types of therapeutic agents.
Sustained Lidocaine Release through Medicated Submicron Particles in Thermo-reversible Smart Gels
.Using biodegradable PC particles, made out of oppositely charged biomacromolecules (e.g. negatively charged alginate and positively charged chitosan), we partially demonstrated that entrapped lidocaine molecules (that have formed a partial complex with alginate) could be released over a relatively longer period of time (up to 10 days as shown in Figure 10). In this section, we want to apply another new pathway to further sustain the release profile, i.e., the use of a recently developed Flash NanoPrecipitation method based on a confined impinging jets (CIJ) mixer to fabricate surfactant free lidocaine-containing micron- and sub-micron-sized particles.
Flash NanoPrecipitation Method to Fabricate Medicated Particles
In the present invention (sub-)micron size particles containing lidocaine molecules in either alginate/chitosan PC, or their mixtures using the recently developed process termed Flash NanoPrecipitation are produced. Although the therapeutic agent used in the present invention is lidocaine, it is within the scope of the invention to use other therapeutic agents in place of or in addition to lidocaine. This process utilizes a confined impinging jets (CIJ) mixer as shown Docket No. 788-74 (R7921) schematically in Figure 13. In the CIJ mixer, two high velocity jets (with speeds up to a few meters/sec) impinge each other through a coUinear aligned capillary (marked light green). The separation length scale of the incoming fluids is rapidly reduced by these two equal momentum collinear jet streams. The size of the mix chamber (marked dark green) is sufficiently large to allow for the formation of an impingement lane, but confined to avoid significant "bypassing", i.e., without encountering the other opposing stream. The characteristic mixing time can be adjusted by changing the speed of the impinging jets. The mixing time D T1,^ can be as short as a few milliseconds. If the jet velocities in the region that have characteristics of a turbulent-like flow pattern within the cylindrical mixer, it will further reduce the mixing time. The reaction products after the rapid mixing will be guided through a long tube runner (marked blue) to ensure that the two opposing steams were fully reacted before final collection.
For fabrication of medicated PC particles, one impinging stream contains cationic chitosan and slightly cationic lidocaine, while the other stream contains anionic alginate. Both PC components (i.e., chitosan and alginate) in the solution will be precipitated out upon mixing, and the weak formed complex (and only slightly insoluble) lidocaine-alginate complex can be entrapped within the particle. The process is based on the supposition that lidocaine will react with alginate and form the complex that can be incorporated into the PC. If the mixing is faster than the induction time for precipitation, the process will be in a "homogeneous" condition, i.e., the effect of mixing is not convoluted with the precipitation times. In other words, if mixing time τmjx is sufficiently short, one does not have to consider the mixing effects. This condition permits only two sub-processes to compete: the precipitation of alginate/chitosan with a characteristic time τaεg and the possibility of lidocaine-alginate complex formation as well as the 'endless chain' formation for lidocaine in solution and in the hydrated crystal formation, as discussed briefly in 3.3.1. The drug "complex formation" will have a characteristic time τ ng. In fact, as long as the mixing time τmjX is less Docket No. 788-74 (R7921) thanτagg or τng, whichever is the shortest, the size of the formed nanoparticles can be designed by carefully choosing the concentration of drug, which tunes the τng, and the properties of PC which determine τagg.
When the two characteristic times are properly tuned to match each other, both the particle size and the amount of drug inclusion can be controlled, without the use of surfactants. The detailed balance of the two competing processes will be determined by drug concentration, kinetics of PC formation and lidocaine complex formation, but the formed particles will be expected to have a narrow size distribution. The present invention takes advantage of the unique properties of lidocaine in the impinging jet approach, by making it less soluble either when forming a complex with alginate or when forming 'endless chains'. By using this technique we can avoid the use of Span 80, to form more uniform sized lidocaine-containing particles, and to have the pathway for easier scale-up productions. Several narrow size distribution micro- and nano-particles can be mixed together to tailor-design the drug release profile.
In Vivo Test of Lidocaine Containing Thermo-reversible Smart Gels
In vivo experiments to investigate the efficacy of lidocaine containing thermo-reversible smart gels using two rat models: the limb-movement test and the well-established tail-flick test can be carried out. The primary objective of this task is to determine the relationship between the efficacy of medicated gel in a period of one week and the appropriate dosage of lidocaine loading. The initial drug concentration of the gel, to be applied to the rat models, will be formulated based on the release profile to simulate what is required for systemic treatment in human patients after surgery. As the local concentration of the drug around the injured site can be very high, we expect that the drug concentration can be significantly lowered in actual applications, but this will be determined by the in vivo tests for a specific target. The detailed descriptions of these tests are 7 Docket No. 788-74 (R7921) illustrated in the Vertebrate Animals section. With multiple independent variables, results can be compiled and analyzed by running a factorial ANOVA test and a post-hoc Turkey test to determine if there are any statistical differences between means of response times or strain measurements for particular groups.
The present invention directed to smart gel formulations has the following advantages.
• AU the materials (matrix/particles) are biocompatible, bioabsorbable or/and biodegradable. The formulation is ideal for local pain-management on injured sites (internal and topical).
• The smart gel formation is simple, flexible and effective. The controlled release rate of thermo- reversible gels can be designed over a very wide range from hours to days. Pluronic block copolymer blends with thermally reversible gelation behavior near body temperature and water resistant characteristics during drug delivery, will be particularly useful for localized deployment (e.g. aerosol spraying, injection and topical coating)
• The particle formation can be achieved without the use of surfactants. The amount of drugs, together with particle size will provide the essential variables needed to tailor design both the amount and the long-term release rate of drugs.
• By mixing the drug-containing particles with the short-term smart gel approach, we have a platform for local drug delivery that can control the amount of drug delivered and the release rate profile from hours to days and up to many weeks. Numerous types of biomaterial polyelectrolyte complexes can be used for the process.
Vertebrate Animals
Two animal models that can be used to test the efficacy of lidocaine-containing smart gels: the dorsum skin defect model (so called the limb-movement test) and the tail-flick test. Both models are approved by the Stony Brook University (IACUC no. 06-1516 and 06-1554, Docket No. 788-74 (R7921) respectively). (1) The dorsum skin defect model can offer a sophisticated method to evaluate the efficacy of the pain-releasing device. The application of strain gauges to the extremities of the rats to quantify movement shall enable us to gather more analytical data. However, because the rats in this model will be restrained in a sling with their extremities hanging down through openings built into the sling, any "wiggling" movement that may not be a direct result of the laser shined upon the defect would be recorded. Establishing a baseline level of movement across all animals may be difficult, at best. (2) The tail-flick test is a commonly used procedure that is well established in the literature in which only the rat tail is exposed. The response time between the application of the external stimulus and the tail movement is recorded. Over the course of one week, the stimulus will be applied at discretionary intervals, and the response times will be correlated to the amount of drugs that has been released from the gel. This all-or-nothing approach is sufficient for data to determine the efficacy of our drug-delivery device.
The Dorsum Skin Defect Model (The Hot-Plate Test)
The Dorsum Skin Defect Model also known as "the hot plate test" can be used to test Sprague Dawley (SD) rats as an animal model of mammalian vertebrates that appear to experience pain in a similar fashion to humans. Initially, rats will be acclimated to sling restraints that position the rat so their paws are just able to touch the floor. This will be done by placing each rat into the sling for several minutes on multiple occasions until they become accustomed to the device as shown by minimal reaction or movement during the time they are in the sling. After acclimation, strain recorders will be attached to the rat legs and to the sling, with baseline movement levels established for each animal. If an animal fails to acclimate to the sling, it will be removed from the study and replaced. The SD rats will then be divided into 3 groups: (1) a sham surgery control, (2) a group receiving surgery and no treatment, and (3) a group receiving surgery and medicated gel. The initial dosage of Lidocaine in the trial membranes during in vivo study will be calculated based on Docket No. 788-74 (R7921) the clinical dosage (100 mg) given to a patient of 70 Kg, which can be viewed as the upper limit. The optimized dosage will be determined through this in vivo study. The surgery will consist of removal of 1 square centimeter piece of skin on the back, just behind the shoulder blades. The rats will be allowed to recover from anesthesia and then placed in the sling and their level of analgesia assessed. This will be done by exposing the surgical site (or skin in the same location for sham controls) to an infrared light source that will heat the area and intensify any discomfort that might be present. The reaction of the rat to this stimulus will be recorded on videotape as well as via strain gauges attached to the rat legs and to the sling. The amount and force of movement as the animal attempts to escape the light source should correlate to the level of discomfort being experienced by the rat. The sling will prevent random locomotion of the animal that may confound interpretation of their escape response.
The Tail-Flick Test
The rat tail-flick test can be adapted from previously established models, in which the level of discomfort experienced by the rats will be indirectly evaluated from physical movements. In our system, Sprague-Dawley rats (200-250g) are trained to remain stationary while being immobilized in a restraint unit (Lomir Biomedical, Inc.). The animals are divided into groups (n=8) according to their prescribed treatment: a control group (no medication) and analgesic groups. The analgesic groups are further divided by the composition of the controlled drug-delivery systems. A predetermined distal portion of their tails are subjected to thermal stimulation by a focused beam of high-intensity light, and the response times to these stimuli are measured; the extent of movement are also recorded by the number of tail flicks, or any other bodily flinches, that occur. The light source is applied for no longer than 30 seconds, or for 10 seconds after escape movements are noted. The analgesic groups that will have the drug medication sprayed on will have the thermal stimulus applied at 30 minutes post-surgery, 45 minutes, 1 hour, 3 hours, 8 hours, 24 hours and Docket No. 788-74 (R7921) daily up to 7 days to determine the longevity of efficacy of the applied controlled drug delivery systems. Terminal blood samples from each animal will be analyzed for indicators of pain and distress.
In summary the present invention is directed to (1) Medicated smart gels, containing at least one therapeutic agent, such as lidocaine, (an anesthetic (model) drug in the present discussion) and mixtures of Pluronic-based block copolymers (poly(ethylene oxide)-ό-poly(propylene oxide)-Z>- poly(ethylene oxide)) in water or buffer solution which can be fine-tuned by copolymer composition and concentration to possess thermo-reversible gelation behavior near the body temperature. Such characteristics shall enable the sustained and controlled release of lidocaine molecules and can facilitate site deployment pathways (e.g. low pressure spraying, injection or topical coating). (2). The dynamic formation of microscopic bioabsorbable medicated particles (micron and submicron sizes) based on polyelectrolyte complex formation (e.g., mixtures based on chitosan and alginate) can provide a sustained release of lidocaine to tens of days. The combination of these two technologies offers a powerful means to prepare effective pain-management media for site-specific treatments as well as treating other diseases which require a steady administration of drugs with a beneficial delay-release.
While the preferred embodiment of the present invention has been illustrated and described in detail, various modifications of, for example, components, materials and parameters, will become apparent to those skilled in the art, and all such modifications and changes are intended to fall within the scope of the claims of the present invention.

Claims

DOCket INO. 788-74 (K7y.∑ljWHAT IS CLAIMED IS:
1. A thermo-reversible gel comprising:
(sub)micron size particles containing lidocaine, prilocaine, bupivacaine, ropivacaine molecules, analogs and mixtures thereof; and alginate/chitosan polyelectrolyte-complex (PC), chitosan polyelectrolyte-complex (PC) or mixtures thereof.
2. A thermo-reversible pluronic hydrogel comprising a series of (E) and (P) triblock copolymers having a general formula of ExPyEx, wherein E is a hydrophilic molecule and P is a molecule capable of hydrogen bonding with water, hydrophilic at low temperatures, and hydrophobic at higher temperatures due to the break down of hydrogen bonds between the P blocks and water.
3. The thermo-reversible pluronic hydrogel of claim 2, wherein x has a value of from about 1 to about 300 and y has a value of from about 1 to about 300.
4. The thermo-reversible pluronic hydrogel of claim 2, wherein x has a value of from about 30 to about 100 and y has a value of from about 30 to about 100.
5. The thermo-reversible pluronic hydrogel of claim 4, wherein E is ethylene oxide and P is propylene oxide.
6. The thermo-reversible pluronic hydrogel of claim 4, wherein x is from about 30 to about 100 and y is from about 30 to about 100.
7. The thermo-reversible pluronic hydrogel of claim 2, wherein E is ethylene oxide and P is propylene oxide.
8. The thermo-reversible pluronic hydrogel of claim 2, further comprising a therapeutic agent. Docket No. 788-74 (R7921)
9. The thermo-reversible pluronic hydrogel of claim 8, wherein the therapeutic agent is any of lidocaine, prilocaine, bupivacaine, ropivacaine molecules, as well as analogs or mixtures thereof.
10. A thermo-reversible pluronic hydrogel blend comprising: a series of (E) and (P) triblock copolymers with the general formula OfExPyEx, wherein E is a hydrophilic molecule and P is a molecule capable of hydrogen bonding with water, hydrophilic at low temperatures, and hydrophobic at higher temperatures due to the break down of hydrogen bonds between the P blocks and water, wherein x is from about 1 to about 300 and y is from about 1 to about 300; a series of (P) and (E) triblock copolymers with the general formula of PxEyPx, wherein- E is a hydrophilic molecule and P is a molecule that is capable of hydrogen bonding with water and therefore is hydrophilic at low temperatures, but becomes hydrophobic at higher temperatures due to the break down of hydrogen bonds between the P blocks and water wherein x is from about 1 to about 300 and y is from about 1 to about 300; and at least one therapeutic agent.
11. The blend of claim 10 wherein said therapeutic agent is any of lidocaine, prilocaine, bupivacaine, ropivacaine molecules, as well as analogs or mixtures thereof.
12. A drug delivery system which comprises a smart gel, and a therapeutic agent incorporated therein, wherein said therapeutic agent is optionally incorporated into (sub)micron sized particles, wherein said smart gel is a thermo-reversible gel.
13. The drug delivery system of claim 12 wherein said therapeutic agent is incorporated into (sub)micron sized particles, wherein said (sub)micron particles are comprised of an alginate/chitosan polyelectrolyte-complex (PC), chitosan polyelectrolyte-complex (PC) or mixtures thereof.
14. The drug delivery system of claim 12 wherein said particle size ranges from about 10 nm to about 1 micron.
15. The drug delivery system of claim 12 wherein the alginate/chitosan ratio is from 100/0 Docket No. 788-74 (R7921)
to 0/100.
16. The drug delivery system of claim 12 wherein said therapeutic agent is any of lidocaine, prilocaine, bupivacaine, ropivacaine molecules, analogs and mixtures thereof.
17. The drug delivery system of claim 12 wherein said thermo-reversible gel is a thermo- reversible pluronic hydrogel comprising a series of (E) and (P) triblock copolymers having a general formula of ExPyEx, wherein E is a hydrophilic molecule and P is a molecule capable of hydrogen bonding with water, hydrophilic at low temperatures, and hydrophobic at higher temperatures due to the break down of hydrogen bonds between the P blocks and water.
18. The drug delivery system of claim 17, wherein x has a value of from about 1 to about 300 and y has a value of from about 1 to about 300, E is ethylene oxide and P is propylene oxide.
19. The drug delivery system of claim 12 wherein said thermo-reversible gel is a thermo reversible pluronic hydrogel blend comprising: a series of (E) and (P) triblock copolymers with the general formula of ExPyEx, wherein E is a hydrophilic molecule and P is a molecule capable of hydrogen bonding with water, hydrophilic at low temperatures, and hydrophobic at higher temperatures due to the break down of hydrogen bonds between the P blocks and water, wherein x is from about 1 to about 300 and y is from about 1 to about 300; a series of (P) and (E) triblock copolymers with the general formula OfPxEyPx, wherein E is a hydrophilic molecule and P is a molecule that is capable of hydrogen bonding with water and therefore is hydrophilic at low temperatures, but becomes hydrophobic at higher temperatures due to the break down of hydrogen bonds between the P blocks and water wherein x is from about 1 to about 300 and y is from about 1 to about 300; or mixtures thereof.
20. The drug delivery system of claim 19 wherein E is ethylene oxide and P is propylene oxide.
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