text:
Non-woven fibrous scaffolds made by electrospinning from the synthetic biodegradable polymer such as, for example, poly(lactic-co-glycolic acid) (PLGA) and natural proteins, such as, for example, gelatin (denatured collagen) and elastin and a method of making thereof.
1. A fiber comprising an electrospun blend of at least one synthetic polymer and at least two natural polymers, wherein none of the polymers is crosslinked, wherein said synthetic polymer is poly (lactide-co-glycolide) (PLGA) and wherein the at least two natural polymers are elastin and gelatin, further wherein said electrospun blend comprises said synthetic polymer, gelatin and elastin at a ratio of 30:8:40.
2. A scaffold made from the fiber of
3. The scaffold of
4. The scaffold of
5. The fiber of
6. A method of making the fiber of
(a) providing a blend of at least one synthetic polymer and at least two natural polymers, wherein none of the polymers is crosslinked, wherein said synthetic polymer is poly (lactide-co-glycolide) (PLGA) and wherein the at least two natural polymers are elastin and gelatin, further wherein said electrospun blend comprises said synthetic polymer, gelatin and elastin at a ratio of 30:8:40; and
(b) electrospinning the blend to form the fiber.
7. A fiber comprising an electrospun blend of at least one synthetic polymer and at least two natural polymers, wherein none of the polymers is crosslinked, wherein said synthetic polymer is poly (lactide-co-glycolide) (PLGA) and wherein the at least two natural polymers are elastin and gelatin, further wherein said electrospun blend comprises said synthetic polymer, gelatin and elastin at a ratio of 20:16:40.
8. A scaffold made from the fiber of
9. The scaffold of
10. The scaffold of
11. The fiber of
12. A method of making the fiber of
(a) providing a blend of at least one synthetic polymer and at least two natural polymers, wherein none of the polymers is crosslinked, wherein said synthetic polymer is poly (lactide-co-glycolide) (PLGA) and wherein the at least two natural polymers are elastin and gelatin, further wherein said electrospun blend comprises said synthetic polymer, gelatin and elastin at a ratio of 20:16:40; and
(b) electrospinning the blend to form the fiber.
1. Field of Invention
This invention relates to engineering of functional three-dimensional (3-D) tissue constructs for the replacement and/or repair of damaged native tissues.
2. Description of Related Art
Engineering of functional three-dimensional (3-D) tissue constructs for the replacement and/or repair of damaged native tissues using cells and scaffolds is one of the ultimate goals of tissue engineering. The following references describe various scaffolds:
Further information regarding scaffolds can be obtained from U.S. Pat. Nos. 6,306,424, 6,616,435, 6,689,166, 6,753,311, 6,790,455, 6,790,528 and 6,821,479 and U.S. Patent Application Publication No. 2002/0090725 to Simpson et al.
Problems with currently available scaffolds include the following: current scaffold need to be crosslinked, mechanical properties are not suitable for multiple applications, and limited cell in-growth.
Accordingly, there is a need in the art to provide scaffolds with better mechanical properties and improved cell in-growth capabilities.
All references cited herein are incorporated herein by reference in their entireties.
The invention provides fibers and nanofibrous biocompatible biomatrices electrospun from a blend of synthetic polymers, such as PLGA and natural proteins, especially containing elastin, can be used as tissue engineering scaffolds and implanted into the body to replace/repair damaged/non-functional tissues. These particular blends provide a unique mix of mechanical and physical properties that facilitates cell penetration and proliferation within the scaffolds without crosslinking.
Inventors have observed that the fibrous scaffolds made from the blend remain intact in complete cell culture medium with serum for over one month without crosslinking. In addition, the fiber-gel structure creates a fibrous matrix that is malleable and allows migration of cells throughout the scaffold.
In this invention, non-woven fibrous scaffolds were electrospun from a tertiary blend of the synthetic biodegradable polymer such as, for example, poly(lactic-co-glycolic acid) (PLGA) and natural proteins, such as, for example, gelatin (denatured collagen) and elastin. The electrospinning process is attractive for fabrication of scaffolds for tissue engineering due to the simplicity of system set-up and ability to generate nanoscale features. In addition, the electrospinning process is capable of generating fibrous scaffolds from both natural and synthetic polymers. The limitation of scaffolds fabricated using the electrospinning process is the high fiber density and the resultant "fish net effect". In other words, fiber density in electrospun mats is often too high to allow for ingrowth of cells. Possible solutions include the use of composite scaffolds containing heterogeneous fibers that display different solubility in a given solvent to create "spacers". The major drawback of fibrous protein scaffolds is the need for crosslinking in order to prevent rapid hydrolysis of the fibers. In order to create composite scaffolds that do not require crosslinking and display enhanced capacity to support the 3-D tissue assembly, the present invention offers a solution to this problem by using a tertiary blend fibers of polymers and natural fibers such as PLGA, gelatin, and elastin.
The initial plan was to use co-spinning (each material in different syringe) of PLGA and natural proteins such as collagen without crosslinking, so that when the scaffold was placed into aqueous medium the non-crosslinked protein fibers would dissolve, producing a porous PLGA fiber mesh. This approach seemed interesting, however, inventors decided to try dissolving multiple materials in the same solution to see if this would produce an even more porous scaffold. To investigate this possibility, mixed spinning (all materials in the same solution) of a tertiary blend of PLGA, gelatin, and elastin was carried out. Unexpectedly, the resultant fiber mat became a solid fiber-gel composite upon placement into aqueous medium, rather than the non-crosslinked components completely dissolving. The PGE mats were visibly swollen in aqueous medium and had a gel-like appearance, although it was clear that there were still solid fiber components in the scaffold. Furthermore, seeding of PGE scaffolds with bone marrow stromal cells resulted in contraction of PGE scaffolds to <50% of the original swollen size.
Electrospinning provides an efficient approach to fabricating biomimetic scaffolds from both natural and synthetic polymer and proteins for tissue engineering. Inventors electrospun a fibrous composite PGE scaffold composed of PLGA, gelatin, and elastin. The advantage of this scaffold was the beneficial characteristics of each individual component synergizing to form a composite with more superior three dimensional cell culture properties. This conclusion is corroborated by histological analyses indicating that rBMSCs attach, migrate, and proliferate within the PGE matrix, resulting in extremely high cell densities by 4 weeks of in vitro culture. These scaffolds would have wide applicability in a number of tissue engineering systems due to their ability to support dense cell growth, thus having the ability to deliver extremely high cell numbers. This novel approach to generating composite scaffolds of natural and synthetic biomaterials affords tissue engineers the ability to meet all necessary design criteria in fabricating scaffolds for a given application.
The invention will be described using poly(lactic-co-glycolic acid)-gelatin-elastin fibers (PGE) as a non-limiting example. In the absence of a cross-linking agent, the average PGE fiber diameter increased from 347�103 nm to 999�123 nm upon wetting as measured by scanning electron microscopy. Rat bone marrow stromal cells (rBMSC) were used paradigmatically to study the 3-D cell culture properties of PGE scaffolds. Consistent with the observed properties of the individual fibers, PGE scaffolds initially swelled in aqueous culture medium, however rBMSC seeded PGE scaffolds contracted to <50% of original size. Time course histological analysis demonstrated uniform seeding of rBMSC into PGE scaffolds and complete cell penetration into the fibrous architecture over 4 weeks of in vitro culture.
Accordingly, the invention includes a fiber comprising an electroprocessed blend of at least one synthetic polymer and at least two natural polymers without crosslinking, wherein one of the at least two natural polymers is elastin or a pro-form of elastin or an elastin-like materials.
In certain embodiments, one of the at least two natural polymers is gelatin.
In certain embodiments, said synthetic polymer is at least one of poly (epsilon-caprolactone) (PCL), poly (lactic acid) (PLA), poly (glycolic acid) (PGA), and the copolymers poly (lactide-co-glycolide) (PLGA) or polyaniline.
In certain embodiments, said electroprocessed blend is an electrospun blend.
In certain embodiments, said electrospun blend comprises said synthetic polymer, denatured collagen and elastin at a ratio of 3:1:2 (v/v/v).
In certain embodiments, said electrospun blend comprises said synthetic polymer, denatured collagen and elastin at a ratio of 2:2:2 (v/v/v).
Another object of the present invention is to provide a scaffold made from the fiber as described above.
In certain embodiments, the scaffold further comprises a cell. In certain embodiments, the scaffold further comprises a pharmaceutical agent.
A further object of the present invention is to provide a method of making a fiber as described above, the method comprising providing a blend of at least one synthetic polymer and at least two natural polymers without crosslinking, wherein one of the at least two natural polymers is elastin or a pro-form of elastin or an elastin-like materials; and electroprocessing the blend to form the fiber. In certain embodiments, said electroprocessing is electrospinning.
Yet another object of the present invention is to provide a method of using a scaffold, the method comprising contacting the scaffold with a cell.
In certain embodiments, the fiber comprises a pharmaceutical agent.
In certain embodiments, collagen is excluded from the fiber.
The invention will be described in conjunction with the following drawings in which like reference numerals designate like elements and wherein:
Extracellular matrix proteins are important modulators of the cellular microenvironment, making them attractive choices for the fabrication of tissue-engineered scaffolds [1-3].
Engineered scaffolds made solely out of these proteins provide some measure of functionality and elasticity [3,12,13]. Fabrication of collagen and elastin scaffolds using the electrospinning process requires chemical cross-linking in order to prevent hydrogel transition and eventual dissolving of the fibers in aqueous medium (28). This chemical cross-linking facilitates mechanical stability of the fibers in aqueous medium, but not without it's drawbacks. The fibers become rigid and resistant to cell penetration and certainly lose some degree of biological activity due to cross-linking of amino acid residues. Collagen hydrogels allow for uniform cell growth throughout the 3-D architecture and maintain integrin mediated adhesions (30). however hydrogel scaffolds lack the mechanical properties required. The mechanical properties of these ECM protein hydrogel scaffolds can be enhanced using synthetic, biodegradable polymers, such as poly(lactic-co-glycolic acid) (PLGA). We hypothesize that electrospinning of natural and synthetic polymers dissolved in the same solution without subsequent cross-linking of the protein components would allow for the formation of a stable fiber-hydrogel composite upon introduction into aqueous medium.
In the present invention, electrospinning is a preferred form of electroprocessing. The term "electroprocessing" shall be defined broadly to include all methods of electrospinning, electrospraying, electroaerosoling, and electrosputtering of materials, combinations of two or more such methods, and any other method wherein materials are streamed, sprayed, sputtered or dripped across an electric field and toward a target. The electroprocessed material can be electroprocessed from one or more grounded reservoirs in the direction of a charged substrate or from charged reservoirs toward a grounded target. "Electrospinning" means a process in which fibers are formed from a solution or melt by streaming an electrically charged solution or melt through an orifice. "Electroaerosoling" means a process in which droplets are formed from a solution or melt by streaming an electrically charged polymer solution or melt through an orifice. The term electroprocessing is not limited to the specific examples set forth herein, and it includes any means of using an electrical field for depositing a material on a target.
Electrospinning is an attractive process for fabricating scaffolds for tissue engineering applications due to the simplicity of the process and the ability to generate microscale and nanoscale features with synthetic and natural polymers [14, 28]. To date, a broad range of polymers has be processed by electrospinning, including polyamides, polylactides, cellulose derivatives, water soluble polymers such as polyethyleneoxide, as well as polymer blends or polymers containing solid nanoparticles or functional small molecules [15]. More recently the electrospinning process has been employed for producing fibrous scaffolds for tissue engineering from both natural and synthetic polymers [16]. Bowland et al. fabricated a three-layered vascular construct by electrospinning collagen and elastin [17]. To date, electrospun fibrous scaffolds have been fabricated with numerous synthetic biodegradable polymers, such as poly(∈-caprolactone) (PCL), poly(lactic acid) (PLA), poly(glycolic acid) (PGA), and the copolymers poly(lactide-co-glycolide) (PLGA) [18-21]. Electrospun scaffolds have been proposed for use in the engineering of bone tissue [22-24] and cardiac grafts [25]. Similarly, poly(L-lactide-co-∈-caprolactone) [P(LLA-CL)] has been electrospun into nanofibrous scaffolds for engineering blood vessel substitutes [26,27].
Any solvent can be used that allows delivery of the material or substance to the orifice, tip of a syringe, or other site from which the material will be electroprocessed. The solvent may be used for dissolving or suspending the material or the substance to be electroprocessed. Solvents useful for dissolving or suspending a material or a substance depend on the material or substance. Electrospinning techniques often require more specific solvent conditions. For example, collagen can be electrodeposited as a solution or suspension in water, 2,2,2-trifluoroethanol, 1,1,1,3,3,3-hexafluoro-2-propanol (also known as hexafluoroisopropanol or HFIP), or combinations thereof. Fibrin monomer can be electrodeposited or electrospun from solvents such as urea, monochloroacetic acid, water, 2,2,2-trifluoroethanol, HFIP, or combinations thereof. Elastin can be electrodeposited as a solution or suspension in water, 2,2,2-trifluoroethanol, isopropanol, HFIP, or combinations thereof, such as isopropanol and water. In one desirable embodiment, elastin is electrospun from a solution of 70% isopropanol and 30% water containing 250 mg/ml of elastin. Other lower order alcohols, especially halogenated alcohols, may be used. Other solvents that may be used or combined with other solvents in electroprocessing natural matrix materials include acetamide, N-methylformamide, N,N-dimethylformamide (DMF), dimethylsulfoxide (DMSO), dimethylacetamide, N-methyl pyrrolidone (NMP), acetic acid, trifluoroacetic acid, ethyl acetate, acetonitrile, trifluoroacetic anhydride, 1,1,1-trifluoroacetone, maleic acid, hexafluoroacetone.
One limitation of electrospun scaffolds fabricated with single materials is the high fiber density and the resultant fishnet effect, with the fiber density in electrospun mats is often too high to allow for ingrowth of cells. A major drawback of fibrous protein scaffolds is the need for crosslinking in order to prevent rapid hydrolysis of the delicate fibers. Possible solutions include the use of composite scaffolds containing heterogeneous fibers of natural and synthetic origin which display differential gel-solid state transition upon introduction to aqueous medium in order to create fiber-hydrogel composites. With that, the solid state synthetic polymer material can function as a fibrous backbone while non-crosslinked protein hydrogel components act as promoters of cellular attachment and growth. In this invention, PLGA, gelatin and elastin were dissolved in the same HFP solvent and electrospun this mix into a 3-D fibrous mat. Upon exposure to aqueous medium the resultant composite PGE scaffold became a stable fiber-laden hydrogel with suitable mechanical properties without need for chemical cross-linking. Furthermore, PGE scaffolds displayed superior capacity to support 3-D tissue-like assembly when compared to fiber scaffolds constructed from the individual components. Thus, this novel composite scaffold exhibited some of the advantages of the individual components, providing proof of concept for the engineering design principles.
Synthetic polymers of the invention include biodegradable and non-biodegradable polymers. Such polymers include but are not limited to the following: poly(urethanes), poly(siloxanes) or silicones, poly(ethylene), poly(vinyl pyrrolidone), poly(2-hydroxy ethyl methacrylate), poly(N-vinyl pyrrolidone), poly(methyl methacrylate), poly(vinyl alcohol), poly(acrylic acid), polyacrylamide, poly(ethylene-co-vinyl acetate), poly(ethylene glycol), poly(methacrylic acid), polylactides (PLA), polyglycolides (PGA), poly(lactide-co-glycolid-es) (PLGA), polyanhydrides, and polyorthoesters or any other similar synthetic polymers that may be developed that are biologically compatible. The term "biologically compatible, synthetic polymers" shall also include copolymers and blends, and any other combinations of the forgoing either together or with other polymers generally. The use of these polymers will depend on given applications and specifications required. A more detailed discussion of these polymers and types of polymers is set forth in Brannon-Peppas, Lisa, "Polymers in Controlled Drug Delivery," Medical Plastics and Biomaterials, November 1997, which is incorporated by reference as if set forth fully herein.
Examples of natural polymers of the invention include, but are not limited, to amino acids, peptides, denatured peptides such as gelatin from denatured collagen, polypeptides, proteins, carbohydrates, lipids, nucleic acids, glycoproteins, minerals, lipoproteins, glycolipids, glycosaminoglycans, and proteoglycans. In certain embodiments, collagen is excluded.
Pharmaceutical agents suitable herein can be organic or inorganic and may be in a solid, semisolid, liquid, or gas phase. Molecules may be present in combinations or mixtures with other molecules, and may be in solution, suspension, or any other form. Examples of classes of molecules that may be used include human or veterinary therapeutics, cosmetics, nutraceuticals, agriculturals such as herbicides, pesticides and fertilizers, vitamins, salts, electrolytes, amino acids, peptides, polypeptides, proteins, carbohydrates, lipids, nucleic acids, glycoproteins, lipoproteins, glycolipids, glycosaminoglycans, proteoglycans, growth factors, hormones, neurotransmitters, pheromones, chalones, prostaglandins, immunoglobulins, monokines and other cytokines, humectants, metals, gases, minerals, plasticizers, ions, electrically and magnetically reactive materials, light sensitive materials, anti-oxidants, molecules that may be metabolized as a source of cellular energy, antigens, and any molecules that can cause a cellular or physiological response. Any combination of molecules can be used, as well as agonists or antagonists of these molecules.
Pharmaceutical agents include any therapeutic molecule including, without limitation, any pharmaceutical substance or drug. Examples of pharmaceuticals include, but are not limited to, anesthetics, hypnotics, sedatives and sleep inducers, antipsychotics, antidepressants, antiallergics, antianginals, antiarthritics, antiasthmatics, antidiabetics, antidiarrheal drugs, anticonvulsants, antigout drugs, antihistamines, antipruritics, emetics, antiemetics, antispasmodics, appetite suppressants, neuroactive substances, neurotransmitter agonists, antagonists, receptor blockers and reuptake modulators, beta-adrenergic blockers, calcium channel blockers, disulfiram and disulfiram-like drugs, muscle relaxants, analgesics, antipyretics, stimulants, anticholinesterase agents, parasympathomimetic agents, hormones, anticoagulants, antithrombotics, thrombolytics, immunoglobulins, immunosuppressants, hormone agonists/antagonists, vitamins, antimicrobial agents, antineoplastics, antacids, digestants, laxatives, cathartics, antiseptics, diuretics, disinfectants, fungicides, ectoparasiticides, antiparasitics, heavy metals, heavy metal antagonists, chelating agents, gases and vapors, alkaloids, salts, ions, autacoids, digitalis, cardiac glycosides, antiarrhythmics, antihypertensives, vasodilators, vasoconstrictors, antimuscarinics, ganglionic stimulating agents, ganglionic blocking agents, neuromuscular blocking agents, adrenergic nerve inhibitors, anti-oxidants, vitamins, cosmetics, anti-inflammatories, wound care products, antithrombogenic agents, antitumoral agents, antiangiogenic agents, anesthetics, antigenic agents, wound healing agents, plant extracts, growth factors, emollients, humectants, rejection/anti-rejection drugs, spermicides, conditioners, antibacterial agents, antifungal agents, antiviral agents, antibiotics, tranquilizers, cholesterol-reducing drugs, antitussives, histamine-blocking drugs, monoamine oxidase inhibitor. All substances listed by the U.S. Pharmacopeia are also included within the substances of the present invention.
Material and Methods
Gelatin (bovine skin, type B powder), alpha-elastin (soluble bovine, lyophilized powder) were purchased from Sigma (St Louis Mo.). Poly(lactide-co-glycolide) (PLGA) 90/10 was synthesized according to and processed into porous foams. In brief: The mixture chloroform solution of PLGA and sugar (10˜15 wt % and 90˜85 wt % respectively) was cast in a glass dish and put it in air for removal of the chloroform, and then it is put into water to dissolve the sugar. The biodegradation of PLGA in 90/10 is about 6 month. Unless noted otherwise all chemicals and reagents, including 1,1,1,3,3,3 Hexafluoro-2-Propanol (HFP) and Chloroform (CHCl3) were purchased from Sigma-Aldrich.
Purifying PLGA
PLGA 90/10 foam was added to in a large amount of chloroform with shaking at RT until fully dissolved. Then the PLGA solution was filtered through a glass microfibre filter (GF/A, circle 55 mm �, Whatman�) with vacuum assistance. Finally, PLGA was dried in the chemical hood by overnight evaporating all chloroform at RT. Final concentration: ˜30 wt %=300 mg/ml in HFP (30% P-H).
Final Solution: Mix approximately 1 ml 30% P-H, 0.7 ml 8.3% gelatin, 1.3 ml 20% elastin together. Stir w/ heating the solution in water bath for about 30 min.
Electrospinning was carried out as follows: polymer materials (PLGA, gelatin and alpha-elastin) were dissolved in HFP at concentrations of 30%, 8% and 20% (w/v), respectively, and then mixed as a blend PGE solution (in HFP) at the ratio of 2:3:5 (v/v/v). The blend solution was loaded into 5 ml plastic syringes (from Becton Dickinson, BD) equipped with blunt 18 gauge needles. The syringes were then placed in a syringe pump (14-831-1, KD Scientific Single-Syringe Infusion Pump, from Fisher) and the needles connected to the positive output of a high voltage power supply (ES30-0.1P, Gamma High Voltage Research, Inc.). Fibers were electrospun onto a grounded brass plate (25 mm�75 mm). The solution delivery rate was set at 1.5 ml/h. Electrospinning was carried out with the high voltage power supply set at 10 kV and an air gap distance of 15 cm.
Scanning Electron Microscopy
For analysis of this PGE constructs, scaffolds were sputter-coated with Au/Pd and examined with a scanning electron microscope (SEM, XL-30 Environmental SEM-FEG). Each micrograph from an SEM scan was digitized. The average fiber diameter was calculated using UTHSCSA ImageTool 3.0 software.
Microtensile Test
The tensile properties of electrospun 3-D fibrous protein scaffolds, approximately 0.2-0.5 mm�5 mm�50 mm (H�W�L), were characterized in a Kawabata Evaluation System (KES-G1, Kato Tech Co., Japan) using routine mechanical testing methods for fabric materials. The specimens were mounted vertically on two 1�1 cm mechanical gripping units of the tensile tester, leaving a 40 mm gauge length for mechanical loading. A strain rate of 0.2 mm/sec, sensitivity of 5�10 (10 V=1000 grams), and frequency of 50 Hz were used in the tensile tests.
Scaffold Sterilization
PGE scaffolds were sterilized overnight by soaking in 10% antibiotic containing phosphate buffer saline (PBS) followed by a short exposure to UV light (10 minutes on each side). The scaffolds were then washed with PBS 3 times vigorously to remove residual antibiotic, followed by several washes with serum-free Dulbecco's Modified Eagle's Medium (DMEM) to condition the surface for cell attachment.
Cell Seeding
The PGE scaffolds (5�5�0.25 mm) were placed in a 100 mm Petri dishes on an orbital shaker and seeded overnight using a 500,000/mL suspension of rat bone marrow stromal cells (rBMSC) in complete cell culture medium (see below).
PGE scaffolds were seeded using a 500 K/mL suspension of bone marrow stromal cells (BMSCs) overnight on the Belly Dancer. As mentioned earlier, the PGE scaffolds contracted to <50% of their size prior to seeding during the ˜16 hr seeding process.
Cell Culture
rBMSCs (a gift of Dr. Itzhak Fischer, Department of Neurobiology and Anatomy, DUCOM) were maintained essentially as previously reported (33). Briefly, a 50/50 mixture of alpha-MEM and Ham's F-12 media containing 10% FBS and antiobiotics was used. The BMSCseeded PGE constructs were cultured under static conditions at 37� C. in a 5% CO2 for up to 28 days, with samples taken at 24 hours, 7 days, and 28 days to evaluate cell in-growth over an extended in vitro culture period.
Extended Culture
BMSCs are maintained in a 50/50 mixture of alpha-MEM and Ham's F-12 media containing 10% FBS and antiobiotics. In these experiments the BMSC-seeded PGE constructs were cultured under static conditions for up to 28 days, with samples taken at 24 hours, 7 days, and 28 days. It would have been desirable to have more time points between 7 and 28 days, however there were limited samples and it was decided to let the samples go for ˜1 month to look at cell ingrowth over an extended in vitro culture period.
Histological Characterization
At 24 hours, 7 days, and 28 days PGE constructs were harvested and fixed in 10% formalin for 1-2 hours at room temperature, followed by overnight fixation at 4� C. The next morning the samples were washed with PBS several times and stored in PBS until embedding and sectioning. The PGE constructs were embedded in OCT compound (Tissue-Tek), frozen at −20� C., and sectioned using the Reichert 1800 cryostat.
Results and Discussion
Electrospun PGE Scaffold
Based on our previous studies on electrospinning parameters [28], PGE was electrospun into mats approximately 0.5 mm thick at 10 kV and 1.5 ml/h flow rate with 15 cm air gap distance. Resultant PGE fibers were homogenous in appearance (
Microtensile Properties of PGE Scaffolds
Inventors compared the mechanical properties of non-cross linked dry fibrous mats prepared from respectively: 8% gelatin, 20% elastin, 30% PLGA, and PGE, using a microtensile tester. As seen in
Cell Growth in PGE Scaffolds
After overnight seeding of rBMSC into PGE scaffolds on an orbital shaker, PGE constructs were transferred to fresh cell culture plate and cultured for a period of 4 weeks. Upon seeding with rBMSC, the PGE scaffolds contracted to <50% of the original swollen size. Histological characterization of the PGE constructs indicated BMSCs uniformly and efficiently penetrated into the scaffolds following the approximately 16 hour seeding period. Histological images of the constructs at 24 hours, 7 days, and 4 weeks postseeding demonstrated that rBMSCs penetrated into the fibrous scaffold after only 24 hours seeding and continued to proliferate over time as evidenced by increasing cell densities in histological sections.
Inventors provided data characterizing the structural and mechanical properties and 3-D cell culture applicability of a novel composite fibrous scaffold composed of PLGA, gelatin, and elastin. PLGA is a gold standard biomaterial that has seen wide applications in tissue engineering and drug delivery confirming superior biocompatibility as compared with most synthetic polymers. For this reason, PLGA was selected as the synthetic component of engineered matrix of the invention.
In addition, inventors believe that the synthetic polymer components of such composite matrices, in this case PLGA, could be used as a delivery vehicle for proteins (21) or DNA (29) as needed to facilitate the molecular events needed for generation of given tissue. Structural proteins such as gelatin and elastin fibrils are an important factor for the mechanical properties of the ECM as well as for or cell attachment and migration. The fibrous morphology of the collagenous and elastic components of the ECM in various organs has been well characterized (6). In an attempt to mimic these properties, our fibrous scaffolds are engineered to be porous, pliable, and elastic, so that cells will be able to integrate and/or push the ECM-like fibers aside as they grow into the scaffolds. This design aim was realized, as evidenced by the formation of a fiber-hydrogel composite with mechanical properties similar to homogeneous elastin that allowed for cell penetration while retaining gross physical properties that would be required for eventual transplantation of constructs generated in vitro.
The finding that PGE matrices have secondary tensile moduli essentially identical to homogeneous elastin is surprising given the fact that the blend contains approx. 2/5 (v/v) of 20% PGLA and 8% gelatin combined and 3/5 (v/v) of 20% elastin. This high contribution of the low tensile modulus of elastin may explain the superior cell growth characteristics, in that the cells can easily penetrate the scaffolds. The swelling and hydrogel formation within the PGE matrix upon exposure to aqueous medium is in our opinion what facilitates the superior tissue-like assembly observed within these matrices after 28 days of in vitro culture the interior of the scaffold contains large cell numbers. These images demonstrate complete penetration of cells with high density throughout the cross-section of an electrospun fiber scaffold. Remarkably, there is no evidence of a necrotic core as is often seen in attempts to engineer macroscopic tissue constructs in vitro, indicating that the porosity of the scaffold might facilitate transport of nutrient into the scaffold and removal of toxic waste. We surmise that the robust cellular penetration observed in the absence of overt necrosis is due to the malleability of the PGE matrix.
This scaffold design is a prototype for future composites that will combine the benefits of engineered synthetic materials and natural extracellular matrix proteins.
The results of these experiments are very exciting in that they demonstrate complete cell penetration of electrospun fiber scaffolds. Such high cell density throughout the cross-section of an electrospun fiber scaffold has not been reported to the best of our knowledge. The major problem with the PGE scaffolds is that we have not been able to reproduce the material properties observed in the first batch. Possible reasons include the use of a different type of PLGA (the PGE described here was fabricated with PLGA (from China) and the more recently fabricated PGE was made with PLGA from Absorbable Polymers, Inc.). PGE fabricated since the original batch do not display the swelling properties upon placement in aqueous medium, nor do they contract when seeded with BMSCs. These scaffolds would have wide applicability in a number of tissue engineering systems due to their ability to support dense cell growth, thus having the ability to deliver extremely high cell numbers.
The invention will be illustrated in more detail with reference to the following Examples, but it should be understood that the present invention is not deemed to be limited thereto.
In this Example, composite scaffolds composed of synthetic and natural materials with physicochemical properties suitable for tissue engineering applications are described. Fibrous scaffolds were co-electrospun from a blend of a synthetic biodegradable polymer (poly-(lactic-co-glycolic acid), PLGA, 10% solution) and two natural proteins, gelatin (denatured collagen, 8% solution) and α-elastin (20% solution) at ratios of 3:1:2 and 2:2:2 (v/v/v). The resulting PLGA-gelatin-elastin (PGE) fibers were homogeneous in appearance with an average diameter of 380�80 nm, which was considerably smaller than fibers made under identical conditions from the starting materials (PLGA: 780�200 nm; gelatin: 447�123 nm; elastin: 1060�170 nm). Upon hydration, PGE fibers swelled to an average fiber diameter of 963�132 nm, but did not disintegrate. Importantly, PGE scaffolds were stable in an aqueous environment without crosslinking and were more elastic than those made of pure elastin fibers. In order to investigate the cytocompatibility of PGE, we cultured H9c2 rat cardiac myoblasts and rat bone marrow stromal cells (BMSCs) on fibrous PGE scaffolds. We found that myoblasts grew equally as well or slightly better on the scaffolds than on tissue-culture plastic. Microscopic evaluation confirmed that myoblasts reached confluence on the scaffold surfaces while simultaneously growing into the scaffolds. Histological characterization of the PGE constructs indicated that BMSCs penetrated into the center of scaffolds and began proliferating shortly after seeding. Our results suggest that fibrous scaffolds made of PGE and similar biomimetic blends of natural and synthetic polymers may be useful for engineering soft tissues, such as heart, lung and blood vessels.
Under these conditions, the solid state synthetic polymer material may function as a fibrous backbone while the associated, non-crosslinked protein components act as promoters of cellular attachment and growth. In this study, we dissolved PLGA, gelatin and elastin in the same solvent (1,1,1,3,3,3 Hexafluoro-2-Propanol, HFP) and electrospun this tertiary blend into fibrous matrices. Upon exposure to aqueous medium the resultant composite PGE scaffold turned into stable fiber-laden hydrogel with suitable mechanical properties without need for chemical crosslinking. Furthermore, PGE scaffolds displayed an enhanced capacity to support 3-D tissue-like assembly when compared to fibrous scaffolds constructed from the individual components. This novel composite scaffold exhibited advantages beyond those of the individual components, providing the proof-of-concept for tailoring scaffolds for specific tissue engineering purposes.
Materials and Methods
Materials: Gelatin (bovine skin, type B powder, catalog no. G9391), alpha-elastin (soluble bovine, lyophilized powder, catalog no. E6527) were purchased from Sigma (St Louis, Mo.). Poly (lactide-co-glycolide) (PLGA) 90/10 was synthesized and characterized at the Changchun Institute of Applied Chemistry, Chinese Academy of Sciences, P. R. China. The biodegradation of 90/10 PLGA is about 6 months. Unless otherwise noted, all chemicals and reagents, including 1,1,1,3,3,3 Hexafluoro-2-Propanol (HFP), were purchased from Sigma-Aldrich.
Electrospinning: Electrospinning was carried out as previously described. 12 In brief: PLGA, gelatin and alpha-elastin were dissolved in HFP at concentrations of 10%, 8% and 20% (w/v), respectively, and then mixed as a blend PGE solution (in HFP) at the ratio of 3:1:2 and 2:2:2 (v/v/v), abbreviated by PGE 312 and PGE 222, respectively. The blend solutions were loaded into 3 ml plastic syringes (from Becton Dickinson, BD) equipped with blunt 18 gauge needles. The syringes were then placed in a syringe pump (14-831-1, KD Scientific Single-Syringe Infusion Pump, from Fisher) and the needles connected to the positive output of a high voltage power supply (ES30-0.1P, Gamma High Voltage Research, Inc.). Fibers were electrospun onto a grounded brass plate (25 mm�75 mm) or 15 mm glass coverslips (Fisher). The solution delivery rate was set at 1.5 ml/h. Electrospinning was carried out with the high voltage power supply set at 10 kV and an air gap distance of 12 cm. For cell culture studies the mats were sterilized by exposure on both sides for one hour each to a conventional ultraviolet (UV) source (30 Watts, 65 cm working distance) in the laminar-flow hood. UV treatment did not result in crosslinking, as assessed by the fact that the fiber morphology in UV-sterilized test mats composed of elastin and/or gelatin fibers was indistinguishable from that of non-treated mats. Moreover, the UV-sterilized elastin and/or gelatin mats dissolved in an aqueous environment, just as did the non-treated ones.
Microtensile Test: The tensile properties of electrospun fibrous PGE and PLGA mats, approximately 0.2-0.5 mm−5 mm�50 mm (H�W�L), were characterized in a Kawabata Evaluation System (KES-G1, Kato Tech Co., Japan) using routine mechanical testing methods for fabric materials. Dry samples were measured after electrospinning and wet samples were measured after soaking samples in ddH2O for 1 hour for thorough wetting. The specimens were mounted vertically on two 1�1 cm mechanical gripping units of the tensile tester, leaving a 15 mm gauge length for mechanical loading. The load cell size was 1 kg (10 Newton), loading resolution was 50 g/V, and the spatial resolution was 0.01 mm. The strain rate used was 0.08/second. In the text, strain (ΔL/L) is a dimensionless parameter, defined as the ratio of the changed length (ΔL) following stress divided by the original length of the sample (L).
Cell Culture and Seeding:
H9c2 rat myoblasts (ATCC CRL-1446; American Type Culture Collection, Rockville, Md., USA) were used for studying cell attachment and proliferation on PGE and PLGA fiber-coated coverslips. H9c2 cells were grown in Dulbecco's Modified Eagle's Medium (DMEM) with 4.5 g/L glucose supplemented with 4 mM L-glutamine, 25 IU/mL penicillin and 25 μg/mL streptomycin (diluted from stock concentrations of 5,000 IU/mLl penicillin and 5,000 ug/mL streptomycin, Cellgro, Herndon, V A), and 10% fetal bovine serum (FBS, Hyclone, Logan, Utah) under standard culture conditions (37� C., 5% CO2), as we previously described.
Neonatal rat bone marrow stromal cells (BMSC), a kind gift from Dr. Itzhak Fischer, Drexel University College of Medicine, were maintained in a 50/50 mixture of alpha-MEM and Ham's F-12 media containing 25 IU/mL penicillin and 25 μg/mL streptomycin, and 10% FBS. We used BMSCs in addition to H9c2 cells in order to (a) validate our findings with more than one cell type and (b) assess the penetration of the cells into the scaffolds by histology (shown below). For the seeding experiments, PGE and PLGA fibers were electrospun as 100-150 μm thick "coatings" on circular 15 mm glass coverslips. These PGE-coated coverslips were sterilized by exposure to UV light for 2 hours on each side before cell seeding, and then secured in 24-well culture plates using Viton O-rings (Cole-Parmer). PGE-coated coverslips were seeded with H9c2 cells as described below in section 2.6. In addition, BMSC were seeded into the 3-D scaffolds, as described below in section 2.7.
Microscopic Analysis: For analyzing the fibrous constructs by SEM, dry scaffolds were sputter-coated with platinum (Pt) and examined with a scanning electron microscope (SEM, XL-30 Environmental SEM-FEG). SEM images of hydrated samples were taken and the fiber dimensions determined after air-drying the samples at room temperature (RT) overnight or, for fast drying, baking them on a hot plate (70-80 degree C., 2-5 minutes). Each micrograph from an SEM scan was digitized. For each sample, the average fiber diameter was calculated from at least 30 randomly selected fibers in SEM micrographs.
For cytological assessment, samples were fixed in 10% buffered formalin for 1 hour at RT and then left overnight in PBS at 4� C. The samples were washed once with 1�PBS, and the cells permeabilized for 15 minutes in 0.2% Triton-X 100 (Sigma) in PBS (Cellgro). Following a gentle wash in PBS, the samples were incubated for 15 minutes in PBS containing 2 μg/mL Hoechst 33258 (Bis-benzimide, Sigma, catalog no. B1155), a nuclear stain, and 1 μg/mL rhodamine-phalloidin (Phalloidin-tetramethylrhodamine B isothiocyanate, Sigma, catalog no. P1951), a specific stain for microfilaments. Samples were visualized on a Leica DMRX upright microscope equipped with the appropriate fluorescence filters. Digital images were acquired using a Leica D C 300FX camera.
To assess the ultrastructure of H9c2 cells cultured on PGE and PLGA fibers, samples were fixed with 2.5% glutaraldehyde for 1 hour at RT and then left overnight in PBS at 4� C. The samples were washed with PBS and then dehydrated in graded ethanol,15%, 30%, 50%, 70%, 85%, 95%, and 100% (twice), for 10 minutes each followed by drying in a critical point dryer (CPD, SPI CPD 7501, West Chester, Pa.), as previously described.36,37 The samples were then sputter-coated with Pt, and visualized by SEM as above.
Attachment and Proliferation Assay: Cell attachment to and proliferation on the scaffolds were tested using H9c2 rat cardiac myoblasts and measured with a continual fluorescence assay (AlamarBlue™ [AB], Biosource, Alameda, Calif.), as previously reported. H9c2 cells were seeded onto the PGE and PLGA fiber-coated coverslips, control glass coverslips and TCPS surfaces at a density of 10,000 cells/well. Following overnight attachment in complete DMEM in a tissue culture incubator, supernatants were removed and 1.5 ml fresh complete DMEM containing 5% (v/v) AB was added into each well. After 4 hours incubation, triplicate 100 μl aliquots of the AB-containing medium were removed from each well for fluorescence measurement. Subsequently the AB-containing medium in the wells was replaced with fresh complete medium without AB. AlamarBlue™ fluorescence in the samples was determined as previously described. 12 For continual assessment of cell proliferation, the AB assay was performed on the same cell population every other day for up to 8 days until the cells reached confluence. The data was normalized for each well to the initial AB fluorescence readings at day 0 (the first AB reading taken 12-18 hours post-plating).
Characterization of Cells on and in 3-D PGE Scaffolds: The PGE scaffolds (5�5�0.25 mm) were placed in a 100 mm Petri dish on an orbital shaker (Belly Dancer™, Stovall, setting 6) and seeded overnight using a 500,000/mL suspension of either H9c2 myoblasts or BMSCs in cell culture medium. As indicated in the results section, samples were harvested at various time points and processed for either scanning electron microcopy (SEM) or routine histology. For SEM the samples were fixed in 2.5% glutaradehyde, first for 1 h at RT followed by overnight fixation at 4 degree C. Samples were processed by dehydration in graded ethanol, followed by critical point drying (CPD), as previously described.36,37 For histological characterization, PGE constructs were harvested at 24 hours and 7 days. The samples were fixed in 10% formalin for 1 h at RT, followed by overnight fixation at 4 degree C. The next morning, fixed PGE constructs were washed with PBS several times and stored in PBS until further processing. The PGE constructs were embedded in Tissue Freezing Medium (Fisher Scientific, catalog no. 15-183-38), frozen at −20� C., and sectioned using the Reichert 1800 cryostat. Cryosections, 10 μm thickness, were stained routinely with hematoxylin and eosin (H&E) as previously described.
Statistical Analysis: Where applicable, all data are expressed as mean +/−standard deviation. Student's t-test and single factor ANOVA were used for parameter estimation and hypothesis testing, with p<0.05 and p<0.01 considered as being statistically significant and extremely significant, respectively.
Results and Discussion
Co-electrospun PGE Fibers
Based on data from the literature and our previous studies defining some critical electrospinning parameters for natural polymers. PGE and PLGA were electrospun into mats approximately 0.2-0.5 mm thick as described in Materials and Methods. Unless specified otherwise, studies were carried out using PGE 312 fibers as the representative blend. Resultant PGE fibers were homogenous, in appearance), and smaller than PLGA fibers. In our experiments, the average fiber diameter of PGE was 380+/−79 nm, which is statistically significantly smaller than pure PLGA fibers, 778+/−202 nm (p<0.01). The diameter of gelatin fibers was 349+/−97 nm and elastin fibers was 605+/−102 nm under similar electrospinning conditions. Bookmark not defined. These results suggest that upon electrospinning a heterogeneous PGE blend yields hybrid fibers composed of a homogenous blend of PLGA, gelatin, and elastin, rather than a mixture of individual PLGA, gelatin, and elastin fibers. This observation is in contrast to previous reports by Boland et al. who report heterogeneous/phase separated deposition of co-spun elastin and collagen fibers. As inferred from the resulting fiber diameters, the size of PGE fibers is significantly smaller than that of elastin and PLGA fibers, and approximately same as that of gelatin fibers, which is remarkable, since in terms of the mass of the ingredients, gelatin is the smallest constituent component.
When placed in cell culture medium, PGE scaffolds, unlike other scaffolds made of natural biopolymers, did not disintegrate, even in the absence of crosslinking. In contrast to the dry fibrous mats the wet PGE scaffolds swelled, resembling a 3-D opaque hydrogel as also confirmed by SEM). The average diameter of soaked PGE fibers increased to 963+/−132 nm (n=30), which is ˜2-3 times larger than the native PGE fibers (p<0.01).
It is well known that PLGA will not swell or disintegrate/hydrolyze during the relatively short (overnight) incubation in an aqueous environment. In these experiments, the appearance of PLGA fibers soaked for up to 8 days was unchanged and remained similar to that of dry fibers (data not shown). PLGA (90/10) used for these studies has a half life of about 6 months and did not degrade significantly during the incubation periods used. Swelling of electrospun PGE fibers and formation of hydrolysis-resistant fiber-hydrogel composites, specifically, of tertiary blends of synthetic biodegradable polymers and natural extracellular matrix proteins has not been reported before. To explain this behavior, we propose that the blend PGE is arranged in such a way that PLGA serves as a backbone with the water-soluble ECM molecules gelatin/elastin arranged facing the aqueous phase adsorbed on it. Current experiments are underway to test this hypothesis.
These results are unique and in contrast to a recent report by Stitzel et al, who also co-electrospun a tertiary blend of synthetic and natural polymers, viz. PLGA, collagen and elastin. Inventors surmise that the differences in our findings may primarily be due to the fact that Stitzel et al. used (a) collagen instead of gelatin and (b) a ratio of synthetic to natural polymers which is quite different from ours.
Prior to hydration, the ultrastructure of the PGE 222 and PGE 312 fibers was quite similar, indicating homogenous fibers for both blends. Upon hydration in either H2O or cell culture medium, both PGE 222 and PGE 312 mats were stable without the need for chemical crosslinking. However, ultrastructural examination by SEM showed somewhat different morphologies. For PGE 222, invenotrs noticed bundling and separation into tiny fibers, around 120�11 nm. In contrast, separation and bundling were less pronounced in PGE 312. The reason for this interesting difference most probably lies in the different volume fractions of PGLA in both blends. The data clearly indicate the crucial importance of choosing the proper ratios of synthetic and natural polymers for custom-tailoring engineered scaffolds. Experiments are underway to elaborate on these observations.
Microtensile Properties of Electrospun PGE Mats
Inventors compared the mechanical properties of non-cross linked fibrous PGE and PLGA mats under both dry and wet conditions, using a microtensile tester. As seen below, the secondary tensile moduli of dry PLGA, PGE 312, and PGE 222 are 504�67 MPa, 254�111 MPa, and 122�23 MPa, respectively. Pure PLGA fibers have a higher tensile modulus than that of either PGE blend; reducing the PLGA contents in the blends also leads to a decrease in the tensile modulus (p<0.05). The elasticity of the fiber mats is affected by the amount of PLGA component in the blend. Pure PLGA produces the most elastic scaffold, while reducing the amount of PLGA in the blend, results in a decrease in the elasticity of PGE fibers (p<0.01). These data suggest that the mechanical properties of pure gelatin and elastin fiber scaffolds can be enhanced by co-spinning with synthetic polymers, such as polylactic/polyglycolic acid (PLA/PGA), to more realistically mimic the mechanical characteristics of natural ECM, as also pointed out by Stizel et al. 14
Comparing the mechanical properties of dry and wet samples, tensile modulus and strain data indicate that wet fibrous mats, made of PGE 312 or PGE 222, are always significantly more elastic than the dry mats. For example, for PGE 312, the tensile modulus decreased from 254�111 MPa (dry) to 43�11 MPa (wet, p<0.01), and the maximum strain increased from 0.04�0.01 to 0.64�0.02 (p<0.01). For PGE 222, the tensile modulus decreased from 122�23 MPa (dry) to 12�4 MPa (wet, p<0.01), and the maximum strain increased from 0.03�0.01 to 0.60�0.09 (p<0.01). These data indicate a ˜20-fold increase in the elasticity of both PGE mats upon hydration. In contrast, hydration did not significantly affect the elasticity of PLGA mats. Measuring tensile properties of wet samples is physiologically more relevant, because when using these mats as tissue engineering scaffolds, cells react with wet fibers in an aqueous environment. These mechanical tests clearly indicate that PGE fibers represent a novel composite material with physicochemical properties unique and dissimilar from those of each of the parental materials.
Cell Proliferation on PGE Mats
To study the in vitro cytocompatibility of the electrospun fibrous scaffolds, PGE 222, PGE 312, and PLGA fibers were electrospun onto 15 mm circular glass coverslips, UV-sterilized and mounted inside the wells of a regular 24-well plate, as described in section 2.4. H9c2 rat cardiac myoblasts were seeded at a density of 10,000 cells/well in triplicate and allowed to proliferate for up to 8 days on PGE 222, PGE 312, PLGA fiber coated coverslips, and tissue-culture treated polystyrene (TCPS) control surfaces. The cells proliferated similarly on all substrates during the 8 day culture period, albeit with some slight differences on day 8. At this time point, there were slightly more cells on the PGE 222 and PGE 312 fiber coated coverslips than on TCPS surface (p<0.05). This is an indication that the 3-D structure of the fibrous substrates facilitates higher cell densities than 2-D TCPS substrates. We conclude that all fibrous substrates are equally cytocompatible for H9c2 cells and support cell proliferation to a degree which is equal to, if not slightly higher than on conventional "gold-standards", such as tissue-culture treated polystyrene.
Morphology of Cells Growing on PGE Fiber-coated Coverslips
In order to assess the morphology and cytoskeletal architecture of H9c2 cells on the different fiber coated coverslips, H9c2 cells were seeded and cultured as described in Materials and Methods.
At36 hours post-seeding, subconfluent H9c2 cells grown on PGE 222 were more spread out than the cells growing on PLGA fibers. Similar observations were also made for cells growing on PGE 312. Under these experimental conditions, H9c2 cells formed confluent monolayers within 8 days on both the PGE and PLGA fibrous substrates, however with some significant morphological differences. On the PLGA substrates, just like on TCPS, the cells formed "conventional", aligned 2-D monolayers, as seen by light microscopy. In contrast, on PGE fiber coated coverslips, the cells seem to form multilayered "sheets", as inferred from the fluorescence micrograph. SEM micrographs confirm this impression, showing an aligned monolayer on PLGA and overlapping of H9c2 cells in the different layers of the fibrous PGE matrix.
Cell Penetration into 3-D PGE Fibrous Scaffolds
In addition to culturing H9c2 cells on PGE fiber coated coverslips, we also investigated the growth and penetration of both cardiac H9c2 myoblasts and bone marrow stromal cells (BMSCs) into thicker 3-D PGE and PLGA fibrous scaffolds (5 mm�5 mm�0.5 mm). The scaffolds were seeded and cultured as described in section 2.4 and 2.6. SEM micrographs, taken after 4 days of culture, confirm that H9c2 cells were able to invade the 3-D PGE fibrous scaffolds while remaining on the surface of 3-D PLGA scaffolds.
For histological analysis, BMSCs were seeded onto the scaffolds on an orbital shaker overnight. BMSCs readily penetrated into the fibrous PGE scaffolds within the first 24 hours, where they continued to proliferate over a 7-day culture period. Both of these figures were taken at the center of the samples demonstrating cell penetration throughout the cross-section of electrospun PGE fiber scaffolds, even as early as 24 hours post-seeding.
We note that both H9c2 and BMSC easily penetrate into the PGE scaffolds. We also note the striking difference in the penetration behavior of H9c2 into PGE and PLGA. This latter observation cannot simply be explained by differences in the mechanical properties of the two kinds of scaffolds. The fiber sizes are quite similar (˜780 nm [PLGA] vs.˜960 nm [PGE]), while PLGA mats are actually more elastic than PGE scaffolds. The latter property would intuitively call for better penetration of the cells into the PLGA scaffold, yet the data show the opposite.
It is believed that the inclusion of natural polymers enhances the interaction of the cells with the scaffolds by engaging cell adhesion molecules such as integrins, which then permit the cells to exert/generate higher traction forces as required for cell motility. Integrin-mediated cell migration in the ECM is essential in tissue development in vivo and cell scaffold interactions in vitro. Another important issue relevant to cell migration into our PGE scaffolds is the ability of the cell to remodel natural ECM proteins for which they possess digestive enzymes. Therefore, we believe the pliability/elasticity of a fibrous ECM, the attachment ligands presented, and the ability of the cells to remodel their environment dynamically interplay to control cell migration in 3-D scaffolds. The interplay between mechanical properties, integrin-mediated adhesion, and ECM turnover in determining cell-scaffold interactions is currently under investigation in many laboratories, including ours.
In tissue engineering, any degradable scaffold is meant to serve as a provisional matrix, i.e., promoting cell adhesion, migration and proliferation until the cells secret and deposit their own ECM. In this context, our PGE blend contains gelatin, rather than type I collagen. Gelatin is frequently used as a substrate for growing endothelial cells, and also as an adhesive substrate that does not alter the native characteristics of stem cells. It is well known, that many of the integrin-binding sites of cell adhesion and differentiation found in collagen are still present in gelatin. In terms of their immunogenic properties, no difference was found in the inflammatory response of collagen or gelatin coated vascular grafts. In keeping with the concept of a provisional matrix, PLGA was chosen as a synthetic biodegradable polymer component that provides enhanced mechanical properties in the short term while degrading over time to allow for cellular invasion with deposition and remodeling of the extracellular matrix.
Structural proteins such as collagen/gelatin and elastin fibrils are an important factor for the mechanical properties of the ECM as well as for cell attachment and migration. The fibrous morphology of the collagenous and elastic components of the ECM in various organs has been well characterized.6 In an attempt to mimic these properties, our fibrous scaffolds are engineered to be pliable and elastic via incorporation of PLGA, while providing natural ECM proteins as adhesion ligands via incorporation of gelatin and elastin, so that cells will be able to integrate with and invade into the scaffolds. This design aim was realized, as evidenced by the formation of a fiber-hydrogel composite with mechanical properties similar to homogeneous elastin12 that allows for cell penetration while retaining gross physical properties that would be required for eventual transplantation of constructs generated in vitro.
Importantly for in vitro and in vivo applications, when placed in an aqueous environment, the fibrous scaffolds remain intact and acquire a high degree of elasticity in the absence of crosslinking agents. The lack of chemical cross-linking also preserves the chemical structure of the natural protein constituents, thus allowing for enhanced cell adhesion/migration as well as ECM turnover. The swelling and hydrogel formation of the PGE matrix upon exposure to aqueous medium is in our opinion what facilitates the rapid tissue-like assembly observed within these matrices.
These scaffolds synergistically combine the beneficial characteristics of each individual component to form a composite which is stable in an aqueous environment and, due to its superior mechanical/elastic properties as well as presentation of natural ECM protein adhesion ligands to the cells, facilitates attachment and migration into the 3-D scaffolds. It is believed that composite scaffolds of synthetic and natural polymers, such as the PGE scaffolds described here, will have wide applicability in a number of applications in engineering soft tissues, such as blood vessels, and cardiac or pulmonary tissues. Hence, this scaffold design may represent a prototype for future composites that will combine the benefits of engineered synthetic materials and natural extracellular matrix proteins.
Cell adhesion, morphology, growth, and proliferation on/in scaffolds mimicking the extracellular matrix are largely mediated by integrins. The subunit expression of integrins will vary among different materials, in a cell-specific manner. Differential integrin expressions of cells cultured on electrospun PGE and the starting materials, PLGA, gelatin, elastin, respectively will be further studied. Specifically, differential integrin expression of cells growing in 3-D scaffolds vs. those of cells grown on 2-D surfaces coated with a thin film will be assessed by western blots and immunohistochemistry.
While the invention has been described in detail and with reference to specific examples thereof, it will be apparent to one skilled in the art that various changes and modifications can be made therein without departing from the spirit and scope thereof.
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A pneumatic nebulizer that produces a high volume of aerosols for inhalent delivery of medications and other constituencies. High pressure gas formed into a gas jet is passed through a thin choked region of fluid that is entrailed and impinged upon an aerosol amplifier which creates a spray whose aerosol components are directed up through vents to an aerosol outlet for delivery. Larger-sized liquid particles are caused to pool up in a region surrounding the aerosol amplifer and then drip down back into the liquid medication reservoir.
1. A nebulizer that converts a solution to an aerosol for inhalation delivery of solute medication or other constituencies, comprising:
(a) a nebulizer housing;
(b) a compressed gas inlet through which a source of pressurized gas enters the nebulizer;
(c) at least one orifice through which compressed gas passes from the gas inlet and forms a gas jet;
(d) a supply of liquid containing solute medication which is entrained into said gas jet to form a spray;
(e) an aerosol amplifier surface upon which the spray is caused to impinge to form an aerosol containing liquid particles of various sizes;
(f) multiple spray posts spaced radially around said amplifier surface and said at least one orifice, which cause significant pooling of liquid between said posts around said amplifier surface but are positioned and spaced so as to prevent pooling of liquid on said amplifier surface; and
(g) said multiple spray posts oriented to allow excess pooling liquid to return to said supply of liquid.
2. A nebulizer made according to
3. A nebulizer that converts a solution to an aerosol for inhalation delivery of solute medication or other constituencies, comprising:
(a) a nebulizer housing;
(b) a compressed gas inlet through which a source of pressurized gas enters the nebulizer;
(c) a primary orifice through which compressed gas passes from the gas inlet and forms a gas jet, said primary orifice having a gas jet exit plane defined by a point at which the gas jet leaves the primary orifice, said gas jet exit plane being perpendicular to the direction of the gas jet;
(d) a secondary orifice aligned axially with said primary orifice, said secondary orifice having a diameter greater than said primary orifice, said secondary orifice having a gas jet inlet plane and a gas jet exit plane;
(e) said primary orifice gas jet exit plane and said secondary orifice gas jet inlet plane being spaced apart to form an inter-orifice gap which is in liquid communication with a supply of liquid containing solute medication, said gas jet creating a partial vacuum in said inter-orifice gap so that said liquid containing solute medication is caused to be entrained into said inter-orifice gap so said liquid is mixed with said gas jet in said secondary orifice to form a spray which exits said secondary orifice through said gas jet exit plane;
(f) an aerosol amplifier positioned so said spray impinges thereon to form an aerosol containing liquid particles of various sizes;
(g) multiple spray posts spaced radially around said amplifier and said secondary orifice, which cause significant pooling of liquid between said posts around said amplifier but are positioned and spaced so as to prevent pooling of liquid on said amplifier;
(h) said multiple spray posts oriented to allow excess pooling liquid to return to said supply of liquid; and
(i) an outer containment baffle positioned peripherally about said multiple spray posts and internal to said nebulizer housing, positioned so as to intercept liquid passing radially outwardly of said multiple spray posts.
This application claims the benefit of U.S. Provisional Application No. 60/140,275, filed Jun. 18, 1999, entitled "High Efficiency Medical Nebulizer".
1. Field of the Invention
This invention pertains generally to aerosol generating devices, and more particularly to a high efficiency pneumatic nebulizer, which may be used to deliver medicinal aerosols.
2. Description of the Background Art
Nebulizers of various types have been widely used for inhalation delivery of aerosols containing medication or other constituents to the conductive airways of the lung and the gas exchange regions of the deep lung. Aerosols are relatively stable suspensions of finely divided droplets or solid particles in a gaseous medium, usually in air or oxygen. When inhaled, aerosol particles may be deposited by contact upon the various surfaces of the respiratory tract leading to potential injury, desirable therapeutic action, or planned diagnostic behavior depending on the particular properties of the particles. Inhalable aerosol particles are those with an aerodynamic equivalent diameter between 1 and 5 micrometers.
Due largely to the high permeability of the lung and the copious blood flow, medications depositing in the lung readily enter the blood for action throughout the body, while other medications can directly influence the airway epithelium and effect responses via various airway receptors. Properly generated and formulated aerosols can therefore be helpful in medical treatment. As tracers of airflow or indicators of lung responses, other types of aerosol particles deposited in the lung can also be a valuable diagnostic tool.
A nebulizer produces aerosol of fine particles by breaking a fluid into fine droplets and dispersing them into a flowing stream of gas. The droplet size from a medical nebulizer is considerably smaller than a conventional spray atomizer. Medical nebulizers are designed to convert aqueous solutions or colloidal suspensions to aerosols of fine, inhalable droplets that enter the lungs of a patient during inhalation and are then deposited on the surface of the respiratory airways. Typical pneumatic (compressed gas) medical nebulizers in current use develop approximately 15 to 30 microliters of aerosol per liter of gas in finely divided droplets with volume or mass median diameters in the respirable range of 2 to 4 micrometers.
Medication intended for aerosolization generally consists of a solute (the medication) mixed into a base solution in which the solute may be dissolved. Predominantly water or saline solutions are used with low solute concentrations, typically ranging from 1.0 to 5.0 mg/mL. The delivery of solute to the patient's airway is the primary purpose of a nebulizer. The delivery of solution is usually of minor therapeutic consequence. The output of a medicinal pneumatic nebulizer, carried by a gas, which powers the nebulizer consists of two principal components: vapor, which is solution in gas form without any solute, and small droplets of solution in liquid state which contain a small amount of medicinal solute.
Previously it was generally assumed in the industry that two pneumatic nebulizers having the same gravimetric output, reservoir solute concentration, and gas flow rate, operating under identical conditions would deliver the same amount of solute or medication. The reasoning behind this assumption was that, at the same temperature, both gas streams exiting each nebulizer would have the same percentage of vapor and small droplets, and that each droplet would have a concentration of solute equal to the concentration of the reservoir. This assumption has recently been shown to be false. Large variances in fact exist in medication delivery efficiency between different nebulizers, with nebulizers currently on the market falling at the low end of the possible efficiency spectrum.
The high percentage of vapor among existing nebulizers is the result of internal loss of large particles, which are typically sprayed onto surfaces within the nebulizer. Aerosol production within a pneumatic nebulizer is not a precisely controlled process. Upon production of respirable aerosol within the nebulizer, there are in addition particles created of various others sizes. Of the liquid entrained by a nebulizer jet, typically 1% or less is converted to respirable aerosol particles, which are able to escape the nebulizer. The remaining 99% assume the form of large aerosol particles which are not able to escape the nebulizer (typically>5 μM) and are sprayed onto the internal surface areas of the nebulizer or eventually fall back into the nebulizer reservoir.
As particles are removed from the air stream of the inside of the nebulizer, solution will evaporate to take its place in the form of vapor. The vapor can come from one of two places: from the wetted surface areas internal to the nebulizer, or from other aerosol particles which are still airborne within the nebulizer. Where the vapor comes from is very important to the performance of the nebulizer. Vapor which comes from other airborne particles within the nebulizer, decreases the size of the particle from which the vapor is emerging, making it more probable that the particle will escape the nebulizer, while at the same time increasing the concentration of solute within the aerosol particle. The result is increased quantity of medicinal solute to the patient. Vapor which is formed by the evaporation of solution deposited on the inside surface of the nebulizer has already deposited its solute onto the nebulizer inside surface resulting in less solute, or medication, being delivered to the patient. Additionally the solution which is evaporated off the inside surfaces of the nebulizer, results in solute adhering to the nebulizer surface where it is lost to the patient permanently.
A typical high performing nebulizer might have an aerosol density exiting the nebulizer of 30 μL of fluid per Liter of gas. In general, the gas used to drive the nebulizer is air or oxygen, each of which are capable of carrying 22 μL of water in vapor form per Liter of gas at standard atmospheric conditions. The difference, 8 μL of fluid per Liter of gas, is the minimum guaranteed amount of aerosol particles that are carrying solute or medication. A large portion of nebulizers currently on the market perform at this low threshold, while the remainder generally operates with only marginal improvement.
Therefore, a need exists for a nebulizer which can provide a high concentration of solute (medication) to the patient by minimizing the vapor which is formed by solution evaporating off the inside walls of the nebulizer and maximizing the amount of vapor which is formed by solution evaporating from other aerosol particles which are still within the nebulizer. The present invention satisfies the need for a high performance nebulizer, while it in addition provides this functionality in a condensed and economically manufactured package.
The present invention generally pertains to a pneumatic nebulizer that is able to deliver a high concentration of medication aerosol for a wide range of flow rates and reservoir sizes. The invention is ideal for delivery of medication, which is being carried in a solution, because it is able to deliver a higher medication to solution ratio than existing nebulizers.
By way of example and not of limitation, the present invention employs a jet of gas flowing at the speed of sound to entrain and shatter a stream of fluid against a hemispherical aerosol amplifier which is proximal to the jet orifice. The sonic jet of gas produces a vacuum with respect to atmospheric pressure causing fluid in the nebulizer reservoir to be drawn into the jet of gas. The jet of gas and entrained fluid mix forming small droplets (10-100 μM) which are traveling at a high rate of speed, due to the sonic jet, and are caused to impact on the aerosol amplifier. The aerosol amplifier may be of a variety of shapes provided that it causes the creation of copious amounts of aerosol particles (0-25 μM). A majority of non-respirable particles are produced which are also too large to escape the nebulizer. After striking the aerosol amplifier, the aerosol stream is still travelling at a high rate of speed. Placed circumferentially around the aerosol amplifier are a number of spray posts which collect the bulk of the spray coming off the aerosol amplifier. The spray posts are wide enough and close enough to each other to cause significant pooling of fluid between them without causing pooling of fluid on the aerosol amplifier. Pooling of liquid between the spray posts is primarily the result of the spray posts being close enough that the surface tension of the liquid tends to fill the gap. The pooled fluid acts as a spray baffle which collects most of the aerosol particles greater than 10 μM. Aerosol particles smaller than 10 μM escape by traveling upward between the aerosol amplifier and the posts. Outside the diameter of the perimeter formed by the posts is a containment baffle. The containment baffle is so shaped and positioned to catch large aerosol and spray which may pass through the posts periodically due to sputtering of the fluid build up around the posts. The result is very small wetted surface area. Fluid build up around the spray posts flows down the secondary and returns directly to the fluid reservoir. Minimizing the wetted surface area within the nebulizer is important because it is directly to proportional to the rate of vapor which is formed from solution evaporating off of the internal geometry of the nebulizer.
The particles in the 0 to 5 μM range, which slipped by the posts and the containment baffle, already have a good chance of escaping the nebulizer. The Particles in the 5 to 10 μM range, which also slipped by the baffling action, provide the means for vapor production while at the same time being shrunk to below 5 μM, which will also enable them to escape the nebulizer to be delivered to the patient. Maximizing the aerosol surface area within the nebulizer is primarily a matter of maximizing the number of particles. It is well-known that a volume of gas can only hold up to a certain maximum limit of aerosol particles. Nearing the maximum limit the occurrence of aerosol particle collisions and particle merging increases so as to effectively create the limit. An ideal particle size and a corresponding optimized cutoff limit for the post geometry can be determined in accordance with this limit that varies dependent on total nebulizer volume and flowrate (residence time), but particle size is typically about 10 μM. Producing such a tremendous amount of aerosol, in the range of 0 to 10 μM, creates a situation wherein the combined surface area of the particles far exceeds the wetted surface area inside the nebulizer such that the vapor exiting the nebulizer is primarily derived from the aerosol particles. The result is a high performance nebulizer that produces a high solute to solution ratio aerosol.
An object of the invention is to provide a pneumatic nebulizer which can deliver high solute to solution ratios.
Another object of the invention is to provide the means to achieve the performance of a large volume nebulizer within the geometry of a small or medium sized nebulizer.
Another object of the invention is to maximize the amount of medication that can be delivered to the patient.
Another object of the invention is to minimize the amount of medication deposited on the inside surfaces of the nebulizer.
Another object of the invention is to baffle out the large aerosol particles directly into fluid flowing into the fluid reservoir.
Another object of the invention is to provide the means to control the particle size of the nebulizer.
Further objects and advantages of the invention will be brought out in the following portions of the specification, wherein, the detailed description is for the purpose of fully disclosing preferred embodiments of the invention without placing limitations thereon.
The invention will be more fully understood by reference to the following drawings which are for illustrative purposes only:
FIG. 1 is a perspective view of a nebulizer according to the invention.
FIG. 2 is a perspective bottom view of the nebulizer of FIG. 1.
FIG. 3 is an elevation view of the nebulizer.
FIG. 4 is a top view of the nebulizer.
FIG. 5 is an exploded cross-section view of the nebulizer.
FIG. 6 is an assembled cross-section view of the nebulizer.
FIG. 7 is a close up cross-section view of the nebulizer orifice region.
FIG. 8 is a close up cross-section view of the nebulizer orifice region shown with fluid and aerosol flows.
Referring more specifically to the drawings, for illustrative purposes the high efficiency nebulizer of the present invention is embodied in the apparatus generally shown in FIG. 1. It will be appreciated that the apparatus may vary as to configuration and as to details of the parts without departing from the basic concepts as disclosed herein.
Referring to FIG. 1 the nebulizer 10 of the present invention includes a solute jar 12 connected to an inhaler cap 14 that includes an aerosol outlet 16. FIG. 2 depicts the underside of the nebulizer 10 from FIG. 1, with a compressed gas inlet 18. FIG. 3 is side view of the nebulizer. FIG. 4 shows a top view of the nebulizer wherein the barrel of the aerosol outlet 16 is clearly seen. A cross-section of the disassembled structure is shown in FIG. 5 according to the section lines of FIG. 4. The solute jar 12 consists of the compressed gas inlet 18 whose tubular gas passage continues through a gas inlet column 20 that terminates in a primary orifice 22 which forms a gas jet having a gas jet exit plane perpendicularly disposed to the gas jet. The compressed gas jet inlet 18 of the preferred embodiment has a diameter of approximately 0.250″ for connection to commonly available compressed gas supply tubing. The solute jar 12 retains the solute in the solute fluid reservoir 24. Solute jar threads 26 are used to engage the inhaler cap 14 for retention during normal operation.
An intermediate section 28 is shown in FIG. 5 for attachment to the gas inlet column of the solute jar 12. The section consists of a secondary orifice 30 at the end of an internal cavity 32 having a gas jet inlet plane and exit plane. On the top of the intermediate section is a dished aerosol amplifier receptacle 34 and an underside containment baffle surface 36 retained atop spray posts 38 on a circular section whose periphery is shaped into a secondary post chamfer 40. Secondary orifice 30 is sized to prevent restriction of the sonic jet flow and to assure that an adequate vacuum is generated. The secondary orifice 30 is typically sized to have a diameter of 2 to 2.5 times greater than primary orifice 22. The length of secondary orifice 30 is typically between 0.020 and 0.050 inches, which has been found to be long enough to allow complete mixing of the sonic jet and the entrained liquid without substantially reducing the speed of the resulting spray. Intermediate section 28 fits onto the gas inlet column 20 of the solute jar 12 by placing the internal cavity 32 over the gas inlet column.
Aerosol inhaler cap 14 is held to the solute jar 12 by engagement of the inhaler cap threads 42 with corresponding solute jar threads 26 of the solute jar. The sonic jet containing solute passing through the primary orifice 22 and secondary orifice 30 is aligned to impinge on a hemispherically shaped aerosol amplifier surface 44. Resultant aerosol passes through aerosol vents 46 into an aerosol passage 48 and up through the aerosol outlet where it may be inhaled.
In the cross-section of FIG. 6 are shown the assembled sections 12, 14, 28. A central dashed line rectangle of FIG. 6 is shown magnified in the view of FIG. 7. Within FIG. 7, a vertical inter-flow passage 50 and a horizontal fluid choke 52 are formed when the intermediate section 28 and gas inlet column 20 of the solute jar 12 are connected. The sonic jet exiting the primary orifice 22 creates a localized vacuum that entrains liquid from the fluid reservoir to be drawn up through the inter-flow passage 50 and fluid choke 52. The fluid choke 52, containing an inter-orifice gap 53, represents the area of maximum constriction for the passing fluid, and this choke section is made small enough so as to not interfere with the vacuum created by the sonic jet exiting primary orifice 22. The thickness of the circular fluid choke 52 region typically ranges from 0.008 to 0.020″, depending on desired rate of liquid entrainment. Assembling the intermediate section 28 over the gas inlet column 20 of the solute jar 12 causes the position of the aerosol amplifier 44 to line up directly over and coincident with the secondary orifice 30, while at the same time creating aerosol passages 54. Aerosol in the 0 to 10 μM range is allowed to escape up away from the back pressure created by spray posts 38 through the aerosol passages 54. Spray posts 38 are spaced radially around secondary orifice 30 such that spray coming off aerosol amplifier 44 pools up and between spray posts 38. Optimal size and distance between spray posts 38 depends on quantity of liquid entrained in jet, airflow rate, and surface tension of liquid. The preferred embodiment of the present invention has a gas flow rate of 3 L/min with four spray posts 38 of 0.100″ width spaced on an inside diameter of 0.400″. For the amount of liquid entrained and the gas flow rate, spray posts 38 are spaced close enough to promote a substantial amount of liquid pooling between them while not being spaced so close as to allow pooling of the liquid onto the aerosol amplifier 44, which, in this embodiment, has a diameter of 0.320″. Spray posts 38 are positioned on secondary post chamfer 40 which is immediately adjacent to the physical exit plane of secondary orifice 30. Spray post chamfer 40 helps facilitate the pooling of 10 liquid about spray post 38 without accumulating liquid on aerosol amplifier 44. The distance from the exit of secondary orifice 30 and aerosol amplifier 44 can range from 0.01011 to 0.100″ depending on speed and shape of jet exiting secondary orifice 30 Containment baffle 36 is so positioned to catch any spray which may escape through spray posts 38 due to sputtering. Containment baffle 36 is shaped so that captured liquid accumulates on the outside diameter where it pools and drips back into the solute fluid reservoir 24.
The primary orifice 22 is the primary means of controlling the pressure-flow relationship of the nebulizer. It is well-known that nebulizers function optimally at the maximum obtainable compressed gas jet velocities. Since most clinical settings have a compressed gas supply which is pressurized to only 50 psig, for nebulizers flows of less than 25 L/min, it is impractical to obtain supersonic flow and a simple orifice with a sonic jet is the optimum design. It is well known from compressible fluid dynamics that any orifice exiting into atmospheric pressures with a supply pressure of 13.7 psig or more will have an exiting jet which reaches sonic velocity, and that the speed may be calculated from the equation (KRT)˝, where K is a gas constant (1.4 for air or oxygen), R is the ideal gas constant, and T is the absolute temperature of the gas. For nebulizers running under normal conditions and having a supply pressure of 13.7 psig or greater, the sonic speed of the jet will be approximately 345 M/sec, regardless of the size of the primary orifice 22. Once a desired compressed gas flow rate has been determined for a specified supply pressure, primary orifice 22 may be sized per the equation Qm=ρAV. Where Qm is the mass flow rate of the gas, ρ is the density of the gas flowing through the primary orifice 22, A is the cross-sectional area of the primary orifice 22, which is the dimension desired, and V is the sonic velocity as calculated above.
Fluid and aerosol flow within the nebulizer are shown within FIG. 8. Compressed gas at high pressure enters the nebulizer at the compressed gas inlet (not shown) and travels into the gas inlet column 20. At the end of the column the gas is converted to a sonic jet of high speed gas flowing from the primary orifice 22. The gas passes through the fluid choke 52 region that contains a thin region of retained solute. Sonic jet flow creates vacuum that entrains the solution to become entrained in the gas, the combined flow passing out of a secondary orifice 30. As fluid is entrained into the gas jet flow, replacement fluid is drawn up through the inter-flow passage 50 from the solute fluid reservoir 24. Jet flow with entrained fluid flows through a secondary orifice and strikes the curved surface of the aerosol amplifier 44. As the fluid entrained within the jet flow strikes the aerosol amplifier if forms a spray. Larger spray particles are ejected out radially and pool at the spray posts 38 forming accumulated fluids 56. Resultant aerosol particles 58, of a size in the range of 0 to 10 μM, pass through the aerosol passages 54 and into the large internal cavity formed by the mating of cap 14 and jar 12. Aerosol particles larger than 5 μM are generally too big to escape the nebulizer without first becoming smaller. It is at this time that the gas carrying the aerosol begins to develop its vapor content.
One aspect of the invention is the recognition that if one were to reduce the amount of liquid deposited on the internal surfaces of nebulizer 2, a greater fraction of the increase in vapor content (primarily water vapor content) of the gas passing through the nebulizer will come from the aerosol particles rather than liquid on the internal surfaces of the nebulizer. The present invention achieves this increase in solute concentration of the resultant aerosol particles by the use of spray posts 38 which, as discussed above, causes the non-aerosolized liquid spray deflected from surface 44 to pool up as accumulated liquid 56 between the posts and drip back down into reservoir 24. Because only a small amount of liquid exists on the internal surfaces of the device, the majority of the vapor content is absorbed from the existing aerosol particles; this reduces the diameters and increases the solute concentration of the resultant aerosol particles. Escaping aerosol and gas travel through aerosol vents 46, up aerosol passage 48, and through the aerosol outlet 16 where it can be delivered to a patient, typically through the use of an attaching mask or other means of patient connection.
The present invention therefore provides a nebulizer which can deliver a high solute to solution ratio aerosol. In this way higher levels of medication can be delivered to a patient with less waste. Although the description above contains many specificities, these should not be construed as limiting the scope of the invention but as merely providing illustrations of some of the presently preferred embodiments of this invention. Thus the scope of this invention should be determined by the patented claims and their legal equivalents.
Any and all patents, applications and printed publications referred to above are incorporated by reference.
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John Lasseter Tests Luigi's Flying Tires at Disney California Adventure Park - YouTube
....and perhaps it does, but it is too difficult to pick up the first time around? The speed thing mystifies me too, because you can see from the video just how fast they can get going.
I just remember when this video came out, everyone was salivating at the opportunity to ride it - with many proclaiming that this may be the "sleeper" hit when it opened. And yet six months (and several problems) later, here it is - and all indications are that it is problematic and a disappointment. What gives?


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April 14-20: Historically, America's most tragedy-filled week
By John Boyd : April 17, 2015 : Updated: April 17, 2015 12:21pm
Photo By Associated Press Composite PhotoA week filled with American tragedies
Americans this week are remembering the anniversaries of the plant explosion in West, Texas (top left), and the bombing of the Boston Marathon (bottom right). However, the span of April 14-20 is one that has been filled with tragedies for Americans throughout history. See these other horrific events that happened during this infamous seven-day span in America.
Photo By Fotosearch/Getty ImagesThe Donner Party Begins Its Trek
April 14, 1846 – The Reed and Donner families leave Missouri headed to California for a new life. Along the way, they get lost in present-day Utah and Nevada. Winter storms beset them, supplies run low and the trip intended to be over by the fall stretches into the winter in the high western mountains. Out of food, some members of the party resort to cannibalism of the already dead to stay alive. Only 48 members of the 87-person party make it to California.
Photo By MATHEW BRADYLincoln Assassinated
April 14, 1865 – President Abraham Lincoln is fatally shot while attending a play at Ford's Theatre in Washington, D.C.. He dies one day later.
Photo By Historic Adobe MuseumBlack Sunday
April 14, 1935 – The infamous Black Sunday storm rolls through Texas and Oklahoma, causing untold damage and leading hundreds of thousands of Dust Bowl residents to relocate.
Photo By MARKO DROBNJAKOVICThe Black Hawk Incident
April 14, 1994 – Two military Black Hawk helicopters (not pictured) are shot down over the streets of Ebil, Iraq. Twenty-six people on board are killed, including military and civilians.
Photo By Wire PhotosTitanic Sinks
April 15, 1912 – The RMS Titanic finally sinks after hitting an iceberg the night before. As many as 1,635 passengers and crew members are killed.
Photo By General Photographic Agency/Getty ImagesThe Great Mississippi Flood
April 15, 1927 – Fifteen inches of rain falls in New Orleans, bringing a Mississippi River already swollen from heavy rains upstream beyond its capacity. River levee systems pop in more than 140 places causing flooding in 10 states and killing 246 people.
Photo By DON EMMERT/Getty ImagesBoston Marathon Bombing
April 15, 2013 – Three people are killed when two pressure cooker bombs go off in a terrorist attack during the Boston Marathon.
Photo By Story Sloane Jr. / For the ChronicleTexas City Explosion
April 16, 1947 – The SS Grandcamp explodes at 9:12 a.m. in Texas City, Texas, killing more than 575 people.
Photo By Steve HelberVirginia Tech Shooting
April 16, 2007 – A gunman kills 32 people then himself on the campus of Virginia Tech in Blacksburg, Va. Another 23 are injured.
Photo By RAUL CORRALES/ASSOCIATED PRESSBay of Pigs Invasion
April 17, 1961 – The Bay of Pigs invasion of Cuba by a CIA-sponsored military group is launched. Three days later, the invaders are defeated. Casualties include 118 deaths, 360 wounded and 1,202 captured.
Photo By Chip Somodevilla/Getty ImagesWest Plant Explosion
April 17, 2013 – Fifteen people are killed and more than 160 are injured following an ammonium nitrate explosion at a fertilizer plant in West, Texas.Complete Coverage: Remembering the West Explosion (HoustonChronicle.com)
The British Invade
April 18, 1775 – British troops arrive in Massachusetts to arrest Samuel Adams and John Hancock. Paul Revere embarks on his famous ride to warn them, 'The British are coming!"
San Francisco Earthquake
April 18, 1906 – More than 3,000 people are killed by a 7.8-magnitude earthquake and resulting fires in San Francisco.
Photo By -/AFP/Getty ImagesU.S. Embassy Bombing in Beirut
April 18, 1983 – Sixty-three people are killed, including 17 Americans, when a suicide bomber in Lebanon drives a van filled with explosives into the American embassy.
Photo By William F. Campbell/Time & Life Pictures/Getty ImageUSS Iowa Explosion
April 19, 1989 – A 16-inch turret on a Navy battleship stationed off Puerto Rico unexpectedly explodes, killing 47 seamen in the ship's gun room and wounding the ship.
Photo By BOB STRONG/AFP/Getty ImagesThe Waco Seige
April 19, 1993 – Four federal agents and six members of the Branch Davidian religious group were killed Feb. 28, 1993 when the ATF attempted to raid a sect compound outside of Waco, Texas. After a 51-day standoff, the government led a final assault on the compound resulting in 76 deaths.
Photo By BILL WAUGH/STFThe Oklahoma City Bombing
April 19, 1995 – More than 160 people were killed and almost 700 more injured when domestic terrorists set off a bomb in the Alfred P. Murrah Federal Building in Oklahoma City. More than 320 buildings nearby also suffered some damage.
Photo By MPI/Getty ImagesThe Ludlow Massacre
April 20, 1914 – Approximately 25 striking coal miners and members of their families were killed in a clash with the Colorado National Guard and Colorado Fuel & Iron Company.
Photo By Anonymous/APThe Columbine School Shooting
April 20, 1999 – Two students enter Columbine High School in a Denver suburb and open fire, killing 12 students, a coach and then themselves.
Photo By -/HOPDDeepwater Horizons (BP) Oil Spill
April 20, 2010 – A wellhead blowout on a BP-operated oil platform kills 11 and dumps 210 million gallons of oil into the Gulf of Mexico.
Photo By Eric Risberg/Associated PressA Week of Victories, As Well
As horrifying of a week April 14-20 has been in American history, the country also has had some of its greatest moments, as well. Here are a few.
April 14: The first abolition society is formed (1775); the first commercial movie house opens in New York City (1894); "The Grapes of Wrath" is published (1939); the Human Genome Project is completed (2003).
April 15: The preliminary articles of peace ending the American Revolutionary War are ratified 1783); Jackie Robinson breaks the color barrier in Major League Baseball (1947); McDonald's opens its first franchised restaurant (1955); the Ford Mustang debuts (1964).
April 16: Boston Garden, America's oldest hockey arena, opens (1910).
April 17: George Lucas begins writing "Star Wars" (1973).
April 18: Paul Revere takes his midnight ride (1775); Yankee Stadium opens in New York (1923).
April 19: Americans earn their first Revolutionary War victory over the British at the Battle of Lexington and Concord (1775); "The Simpsons" premieres on as a cartoon short on "The Tracey Ullman Show" (1987).
April 20: The Civil Rights Act of 1871 becomes law (1871).
Americans this week are remembering the anniversaries of the plant explosion in West, Texas and the bombing of the Boston Marathon.
However, the span of April 14-20 is one that has been filled with tragedies throughout American history.
See these other horrific events that happened during this infamous seven-day span in America.


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