Patent Application
20110190679
The present invention relates to an extracorporeal therapeutic device for delivering therapeutic molecules into a body. More particularly, this invention relates to an extracorporeal therapeutic device containing viable cells. One way that the molecules can be delivered into a body is through blood circulation or other bodily fluids. When the device is introduced into the blood stream outside the body, the cells produce and secrete the therapeutic molecule into blood or fluid circulating past the device.
Drug delivery devices useful for introducing therapeutic molecules into a mammal have been the subject of considerable research. In one aspect, the research has focused on the development of devices that deliver molecules produced from cellular metabolism. Efforts have also focused on producing an implantable cell based delivery system that can remain in a patient for an extended period of time.
An implantable device for delivering a pre-selected molecule, for example, a hormone, into a mammal's systemic circulation is described in U.S. Pat. No. 6,716,208, the entire contents of which is incorporated by reference. The device described comprises an implantable blood permeable element that can be anchored to an inner wall of an intact blood vessel and a capsule that is held in place within the blood vessel by the anchored blood permeable element. The capsule encloses viable cells which produce and secrete the preselected molecule into blood passing the capsule. The patent also describes a method for percutaneously introducing the device into a preselected blood vessel.
Intracorporeal cell based delivery devices must be sized to fit within a body, typically a body lumen (such as a blood vessel) and, accordingly, have certain size limitations because of the reduced-size requirements. Further, when an intracorporeal cell based delivery system is designed, the anchors used to attach the device to the body, e.g., blood vessel, must be configured to secure the device without introducing trauma to the body. Finally, implantable devices can be difficult to retrieve, especially if they are left within the body for an extended period of time.
In vitro experiments with a larger extracorporeal device utilizing porcine or human proximal tubule progenitor cells have shown differentiated transport and metabolic functions of the renal tubule assist device (RAD).[references 4-6] The bioartificial kidney (BAK) set-up consists of a filtration unit (a conventional synthetic hemofiltration cartridge) followed in series by the tubule (RAD) unit. The tubule unit is able to maintain viability because oxygen along with metabolic substrates and growth substances are delivered to the tubule cells from both intraluminal ultrafiltrate and blood in the extracapillary space Immunoprotection of the cells is achieved due to the impenetrability of immunoglobulins and immunocompetent cells across the hollow fibers. Rejection of non-autologous cells does not occur.
Pre-clinical studies in large animals have demonstrated that the BAK successfully replaced filtration, transport, metabolic, and endocrinologic functions of the kidney in acutely uremic dogs.[5] Further pre-clinical experiments in acutely uremic dogs have also evaluated the influence of the RAD under stress states. Acutely nephrectomized animals were challenged with infusions of endotoxin (lipopolysaccharide) intravenously or with intraperitoneal administration of doses of viable E. Coli before treatment with either cell or sham control RADs in a BAK.[references 7,8] In these experiments, cell RADs provided metabolic renal replacement and resulted in higher anti-inflammatory plasma levels, better hemodynamic stability, and, in the E. Coli sepsis model, longer survival times compared to sham controls. To further evaluate the role of the BAK in septic shock, a swine model with normal kidney function was given large doses of E. Coli intraperitoneally.[reference 9] All animals developed acute tubular necrosis (ATN) with oligo/anuria within 2-4 hours following administration, and RAD treatment resulted in better cardiovascular performance, lower plasma levels of the pro-inflammatory cytokines, and longer survival times compared to sham controls.
These supportive pre-clinical experiments were the basis for testing human cell RADs in Phase I/II and Phase II clinical trials in intensive care unit (ICU) patients with acute renal failure (ARF) and multi-organ failure (MOF). A favorable Phase I/II safety trial [reference 10] led to an FDA-approved, randomized, controlled, open-label Phase II investigation at 12 clinical sites to determine whether this cell therapy approach alters patient mortality. This Phase II study involved 58 patients, of whom 40 were randomized to RAD therapy and 18 made up a control group with comparable demographics and severity of illness. The early results have been as compelling as the Phase I/II results. Renal cell therapy improved the 28-day mortality rate from 61% in the conventional hemofiltration-treated control group to 34% in the RAD-treated group.[references 11,12] This survival impact continued through the 90- and 180-day follow-up periods (p<0.04), with the Cox proportional hazard ratio indicating that the risk of death was 50% of that observed in the conventional continuous renal replacement therapy group. This survival advantage with renal cell therapy was observed for various etiologies of ARF and regardless of organ failure number (1 to 5+) or the presence of sepsis. Subset analysis of patients with concomitant severe sepsis or septic shock demonstrated an incidence of sepsis of 73% and 67% in the cell therapy and conventional therapy groups, respectively. RAD therapy was associated with a mortality rate of 34% in patients with sepsis, compared to 67% in the conventional treatment group. Thus, these clinical results suggest a major effect on survival rates in these desperately ill patients. The clinical use of renal tubule cell therapy for patients with severe sepsis will not require this complex two-cartridge system with two extracorporeal pump systems, since most are not in ARF.
The present invention provides an extracorporeal cell based delivery system that is designed to address aspects of an intracorporeal cell based delivery system. Further, the present invention provides an extracorporeal device designed to introduce therapeutic agents into a mammal that secretes a pre-selected molecule or a combination of cell products directly into the blood stream or into a body fluid or body cavity and addresses the challenges of the prior art. The invention will be more clearly understood from the description, which follows.
Further, a miniaturized cell therapy device will not require extensive extracorporeal blood pump systems. In addition, a miniaturized device that could be stored at the clinical site for immediate use is required to succeed as a commercial product. The current RAD is stored at a central manufacturing facility at 37° C. and must be shipped at 37° C. to the clinical site, delaying treatment and adding to the cost of therapy. Development of a cell device that can be cryopreserved and stored at clinical sites can help safely bring the device to market.
The present invention provides an extracorporeal therapeutic device for delivery of a pre-selected molecule or cell products into a mammal, for example, into the circulatory system or body fluids of a mammal. An embodiment of the invention enables molecules to be introduced into the circulatory system or a body cavity without invasive surgical procedures. Once the device is deployed, it delivers the molecule directly into the blood stream or body fluid. In addition, the device of the invention is adapted to produce and thereafter secrete the pre-selected molecule or cell product into the blood stream or body fluid over a determined period of time. The extracorporeal device and method provide an easy and reproducible system for delivering therapeutically effective amounts of a gene product, for example, a hormone, growth factor, anti-coagulant, immunomodulator, or the like, directly into the blood stream or body fluid of the recipient without the disadvantages of an invasive implantation procedure.
An extracorporeal device that administers a preselected molecule(s) into the mammal over a predetermined period presents advantages over the prior art. An extracorporeal device has the advantage of being easily taken out of the circulation system compared to the efforts required to remove an implanted device. Accordingly, the present invention provides an extracorporeal device for delivering, over a determined period of time, a preselected molecule or cell products into the systemic circulation of a mammal. In another aspect, the present invention provides a method for non-surgically introducing the device into blood circulation of a mammal that is capable of delivering the preselected molecule or cell products into systemic circulation.
In one embodiment, the device includes a capsule that contains viable cells which produce and secrete the preselected molecule into the blood stream. The device may include an anchoring element, which anchors the device to an inner surface of a tube that circulates blood extracorporeally.
The term “extracorporeal circuit” as used in this specification embraces any tube or conduit outside the body that may be connected to the circulatory system or body fluid compartment in a mammal and provides for the flow of blood or fluid through the tube or conduit by natural (e.g., heart) or artificial (e.g., mechanical pump) circulation. An extracorporeal device of the present invention is configured to be disposed in an extracorporeal circuit.
The term “anchoring element” as used in this specification embraces any structure that may be inserted into the lumen of an extracorporeal circulatory system blood tube or conduit and that, once inserted, may be anchored, for example, by hooks, barbs, or stents, to an inner surface of the tube or conduit. In an exemplary embodiment, the anchoring element may be a blood clot filter-type structure. A variety of blood clot anti-migration filters useful in the practice of the invention are known in the art. The currently preferred anchoring element is an anti-migration filter known as a “Greenfield® vena cava filter”. Useful Greenfield® vena cava filters are described in detail in U.S. Pat. Nos. 4,817,600 and 5,059,205, the entire disclosures of which are incorporated by reference.
The term “capsule” as used in this specification embraces any hollow structure dimensioned to fit within the lumen of a tube or conduit used in an extracorporeal circuit and does not occlude or prevent blood or fluid flow. In one embodiment, the capsule is held in place within the extracorporeal blood circuit by anchoring element(s). For example, the capsule may be retained upstream of the anchoring element, alternatively, the anchoring element may be located downstream of the anchoring element and retained in place by an attachment, for example, a hook or tether, extending from the anchoring element to the capsule. In addition, the capsule may be conical or wedge-like in shape to decrease the turbulence of blood flowing past the capsule. In a preferred embodiment, the capsule is formed from a material that can filter particles such that particles (including cells) below a certain size can pass through and particles above a certain size are prevented from passing through. The filter forms an ultrafiltrate from the blood to minimize the entry of proteins greater than 100,000 molecular weight so that immunoglobulins can be excluded from the bathing media around the cells, especially for nonautologous cells not to activate an immunologic response.
The capsule may include either a single hollow fiber or a bundle of hollow fibers made from a semi-permeable membrane. The semi-permeable membrane filter preferably has pores of a size sufficient to permit the diffusion of a preselected molecule or cell products therethrough but yet small enough to exclude the passage of cells therethrough. The pores preferably are designed to permit the preselected molecule produced by the cells to diffuse directly into the blood stream passing the hollow fiber while preventing the cells from migrating out of the hollow fiber and into the systemic circulation.
A variety of polymers are useful in producing the biocompatible semi-permeable membrane of the present invention. They include, but are not limited to polyalginate, polyvinylchloride, polyvinylidene fluoride, polyurethane isocyanate, cellulose acetate, cellulose diacetate, cellulose triacetate, cellulose nitrate, polysulfone, polystyrene, polyurethane, polyvinyl alcohol, polyacrylonitrile, polyamide, polymethylmethacrylate, polytetafluoroethylene, and polyethylene oxide. In addition, useful semi-permeable membranes may be produced from a combination of such polymers.
In an embodiment, the viable cells enclosed within the semi-permeable hollow fiber(s) of the capsule, preferably are eukaryotic cells, and most preferably are mammalian cells. Although the device described herein may comprise cells which naturally produce and secrete the preselected molecule or other cell products, it is contemplated that genetically engineered cells, i.e., cells transfected with, and capable of expressing a nucleic acid encoding the pre-selected molecule, may likewise be used in the practice of the invention.
In another embodiment, the preselected molecule can be a protein, and most preferably is a hormone, for example, erythropoietin or insulin. It is contemplated, however, that the device may be used to deliver into the systemic circulation any molecule that can be produced and secreted from a viable cell. Although single cell types that produce and secrete a single preselected molecule may be used in the invention, it is understood that cells belonging to a particular cell type that produce and secrete a plurality of preselected molecules likewise may be used in the practice of the present invention. Similarly, it is contemplated that a plurality of cell types, wherein cells belonging to each cell type produce and secrete different preselected molecules, may be combined in a capsule thereby to produce a device that delivers a desirable combination of preselected molecules into the circulation.
Preferred embodiments of the device include three configurations. Each preferred configuration isolates the therapeutic cells to minimize the immune response. In a first configuration, a device consists of a cartridge, a cell bearing unit which may be in the form of tubes attached to the cartridge and an anchoring system. The therapeutic cells are disposed within the tubes and the cells are isolated by the size of the pores in the tube. In a second configuration, the cell bearing unit is in the form of disks that are disposed in the cartridge. The therapeutic cells on the disks are protected from immunologic rejection by isolating the disks in the cartridge and providing pores in the cartridge that prevents the cells from being exposed to undesirable elements while allowing free physiologic exchange for the cells within the extracorporeal blood or fluid stream. A third configuration is a combination of the configurations described above. Other configurations are possible.
The devices utilizing cells disposed in hollow fibers and/or cells disposed upon a substrate, such as disks, also can be included in extracorporeal circuits utilizing pumps and fluid separation devices (e.g., hemofilters, dialyzers, and the like) which can provide the devices with fluids for treatment, such as ultrafiltrate derived from blood or peritoneal fluid or peritoneal fluid itself, such that cells are protected from molecules in the body fluids having the ability to mount an immune response. Such extracorporeal circuits can be used in various applications, including, for example, chronic applications and acute applications. In chronic applications, cells can be isolated from a biopsy taken from the subject to be treated and these cells can be expanded prior to use. Because such cells are derived from a host subject, an immunoprotective barrier, such as a filter, may not be needed. In acute applications, the filters may provide immunoprotection by limiting the immune-stimulating molecules from passing through and contacting the cells. Additionally, the animal is protected from the cells which could elicit an immune response by contact of the molecules produced by the cells. Any cells can be used in embodiments according to the invention, such as eukaryotic cells, such as mammalian cells, and such as renal cells.
Thus, in one aspect of the invention, an extracorporeal therapeutic system includes a housing defining an interior space, a frozen substrate having a trabecular structure disposed within the housing, and at least one cell disposed on the substrate. This aspect or any of the following aspects can have any of the following features. The housing can include an inlet for receiving a fluid and an outlet for releasing a processed fluid. At least one flow separator can be disposed between the inlet and the substrate. The substrate can be initially separate from the housing, and/or can be coated with a biomatrix material such as collagen IV.
In another aspect of the invention, an extracorporeal therapeutic system includes a housing defining an interior space, a substrate including a carbon material coated with niobium disposed within the housing, and at least one cell disposed on the substrate. This aspect or any of the following aspects can have any of the following features. The substrate can have a trabecular structure, and the substrate can be initially separate from the housing. The substrate can be coated with collagen IV. The substrate can also be coated with any three-dimensional biomatrix material that protects the integrity and durability of the carbon material, and the biomatrix material can promote or increase cell expansion, attachment, and/or viability. The housing can include an inlet for receiving a fluid and an outlet for releasing a processed fluid. At least one scaffold to retain the substrate can be disposed within the housing, and at least one flow separator can be disposed between the inlet and the substrate. At least one baffle can also be disposed between the inlet and the substrate. The substrate can be frozen.
In another aspect of the invention, an extracorporeal therapeutic system includes a housing, at least one hollow fiber associated with the housing, a biomatrix material disposed within the at least one hollow fiber, and at least one cell disposed within the biomatrix material. This aspect or any of the following aspects can have any of the following features. The cell can be disposed upon a particle. The biomatrix material can be selected from, but is not limited to, alginate solution, gelled alginate solution, or physiologic buffer, and the biomatrix material can also be selected from nutragen or matrigel. Other examples of biomatrix material can include collagen I or any three-dimensional forming or formed biocompatible matrix. The hollow fiber can be disposed within the housing, and the housing can include an inlet and an outlet.
In another aspect of the invention, a substrate for maintaining cells includes a carbon material coated with niobium and with collagen IV and has a trabecular structure. This aspect or any of the following aspects can have the following features. The substrate can also include a cell that has the potential to provide therapeutic value, such as a eukaryotic cell, a mammalian cell, and/or a renal cell. The substrate can be frozen.
In another aspect of the invention, a method for loading an extracorporeal therapeutic system includes thawing a cryopreserved cell and loading the cell into the system. This aspect or any of the following aspects can have any of the following features. The system can include a housing and at least one hollow fiber associated with the housing, and the cryopreserved cell is thawed, mixed with a biomatrix material, and loaded into the hollow fiber. The system also or alternatively can include the cryopreserved cell disposed on a substrate, which can be loaded into the housing. The substrate can include a carbon material coated with niobium and, optionally, collagen IV.
In another aspect of the invention, a method for extracorporeal therapy includes connecting an extracorporeal therapeutic system to the peritoneum of a mammal, so that a peritoneal fluid of the mammal circulates through the system. This aspect or any of the following aspects can have any of the following features. The system can include a housing and at least one hollow fiber associated with the housing, and the system also can include a biomatrix material disposed within the hollow fiber and at least one cell disposed within the biomatrix material. The system also or alternatively can include a housing and a substrate including a carbon material coated with niobium and, optionally, collagen IV, and at least one cell can be disposed on the substrate. The system also or alternatively can include a housing and a frozen substrate having a trabeculated structure. At least one cell can be disposed on the substrate. The housing can also include an inlet for receiving fluid and an outlet for releasing a processed fluid.
In another aspect of the invention, an extracorporeal cell based therapeutic device includes (a) an anchor system which can be capable of attaching the device to an extracorporeal tube, which when attached to the inner wall of the tube permits blood in the tube to pass therethrough; and (b) a capsule including a plurality of pores and having viable cells disposed therein, so that the capsule, when introduced into the tube, can be retained within the tube by the anchor system and the pores permit nutrients to enter the capsule to maintain viability of the cells disposed therein. This aspect or any of the following aspects can have any of the following features. The capsule can be defined by a semi-permeable membrane, and the semi-permeable membrane can include a material selected from polyvinylidene fluoride, polyvinylchloride, polyurethane, polyalginate, polystyrene, polyurethane, polyvinyl alcohol, polyacrylonitrile, polyamide, polymethylmethacrylate, polyethylene oxide, polytetrafluorethylene, isocyanate, cellulose acetate, cellulose diacetate, cellulose triacetate, cellulose nitrate, polysulfone, and mixtures thereof. The viable cells can be disposed on a plurality of filaments, and the filaments can be metallic. The pores can be dimensioned to prevent passage of antibodies therethrough, and the cells can be disposed on at least one disk. The cells can be eukaryotic cells, such as mammalian cells. The capsule can be adapted to be separated from the tube by attaching and detaching the anchor system from the tube. The capsule can include at least one hollow fiber. The pores can permit solutes less than 150 kD to pass therethrough, and the device can be configured to provide a therapeutically significant amount of a molecule with or without using an artificial blood pump.
In another aspect of the invention, an extracorporeal cell based therapeutic device includes (a) an attachment system anchorable to an inner wall of an extracorporeal tube adapted to provide a conduit for blood, which when anchored to the inner wall of a tube permits blood to circulate; and (b) a capsule defining a plurality of pores and having viable cells disposed therein, so that the capsule can be retained within the tube by the attachment system and the pores permit nutrients to enter the capsule to maintain viability of the cells disposed therein. This aspect or any of the following aspects can have any of the following features. The pores can be dimensioned to exclude an agent in the blood which is detrimental to cell viability. The anchor can be metallic. The capsule when introduced into the tube can be retained upstream of the attachment system, and the capsule can include at least one disk. The cells can also be genetically engineered cells.
In another aspect of the invention, a method of introducing an extracorporeal cell-based therapeutic device into the circulatory system of a mammal includes the steps of: (a) anchoring an attachment system to an inner wall of an extracorporeal tube adapted to provide a conduit for bodily fluid, which when anchored to the inner wall of a tube permits the bodily fluid to circulate; and (b) inserting a capsule defining a plurality of pores and having viable cells disposed therein, so that the capsule can retained within the tube by the attachment system and the pores permit nutrients to enter the capsule to maintain viability of the cells disposed therein. This aspect or any of the following aspects can have any of the following features. In step (a), the attachment system can be introduced into the tube via a catheter, and in step (b), the capsule can be introduced into the tube via a catheter. The device can be used to deliver a therapeutic amount of a molecule in the conduit with or without the use of an artificial body fluid pump. The conduit can be attached to the peritoneum and the peritoneal fluid passes through the conduit.
In another aspect of the invention, an extracorporeal cell based therapeutic device includes (a) an attachment system anchorable to an inner wall of an extracorporeal tube adapted to provide a conduit for a bodily fluid, which when anchored to the inner wall of a tube permits the bodily fluid to circulate; and (b) a capsule defining a plurality of pores and having viable cells disposed therein, so that the capsule, when introduced into the tube, can be retained within the tube by the attachment system and the pores permit nutrients to enter the capsule to maintain viability of the cells disposed therein. This aspect can have any of the following features. The tube can be adapted to be connected to the peritoneum and the bodily fluid includes peritoneal fluid, and the tube can be adapted to be connected to blood vessels and the bodily fluid includes blood. The attachment system can be separable from the tube.
The foregoing and other objects, aspects, features, and advantages of the invention will become more apparent and may be better understood by referring to the following description taken in conjunction with the accompanying drawings, in which:
The present invention provides an extracorporeal device for delivering molecules into a mammal, for example, the systemic circulation of a mammal. The device of the invention is adapted for introduction into an extracorporeal circuit, such as into an extracorporeal blood conduit or a peritoneal fluid circuit. After introduction into the extracorporeal circuit, such as the extracorporeal blood conduit, the device permits the pre-selected molecule(s) or cell product(s) to diffuse out of the device. In the case of the extracorporeal circuit, these pre-selected molecule(s) or cell product(s) diffuse out of the device and into the blood stream or body cavity of the recipient, which in certain aspects does so in response to blood parameters, for example, oxygen tension in the case of erythropoietin-producing cells. Also, the glucose concentrations in the bathing media around the insulin producing cells can stimulate the production of insulin.
The embodiments of the invention have at least some of the following advantages. Because the device is extracorporeal, there is less size constraint and the cell bearing units can be sized larger than if a device was implanted. Further, because the device is not implanted in the body, the anchoring system does not need to be designed to be atraumatic. A cartridge embodiment that fits along a conduit is contemplated without an anchoring system. Because the device is not intended to be inserted in a blood vessel, the introduction of the device to the blood stream is much simpler.
The device includes a capsule or housing that holds cell-bearing material. The cells on the cell bearing material are prevented from provoking an immunological response by isolating the cells using a porous material that creates a barrier between the blood and the cells. The barrier allows the cells to have sufficient physiological exchange (e.g., drawing sustaining nutrients and oxygen) with the blood and delivery of the metabolic products produced by the cells. The metabolic products may be in response to small messenger molecules that might be circulated in the blood stream as a result of the patient's pathologic condition. Also, the metabolic production can be stimulated by introducing a messenger molecule into the bloodstream of the patient. Additionally or alternatively, cells in devices according to the invention are also spared from opportunities of clotting. When a bloodless circuit is utilized, the primary concern of clotting is alleviated by removing contact of the device with blood. Additional concerns related to blood viscosity, nutrient concentration, and immunological response are also alleviated by allowing an altered bodily fluid, such as ultrafiltrate derived from blood or peritoneal fluid, or with peritoneal fluid itself, to contact the cells. In this case, the cells may draw nutrients and oxygen and deliver their metabolic products to these fluids. In one embodiment, with respect to, but not limited to, the chronic application of the device, an autologous cell source can be utilized. When an autologous cell source is utilized, filters placed pre- and post-device in an extracorporeal circuit optionally may or may not be used.
An embodiment of the device of the present invention includes an anchor system that secures the device to an inner wall of a extracorporeal blood circuit. The anchor system may have arms that are adapted to be in a reduced profile configuration during delivery into the blood circuit and expand into a delivered profile so that the arms expand and extend to the blood conduit. The anchoring system may use hooks, barbs, or stents disposed upon the arm of the anchor system. The anchoring system is designed such that when anchored to the wall of the blood conduit, the system permits blood in the vessel to pass through and around the device.
The device includes a shape, e.g., cone, facing upstream that minimizes turbulence of blood as it flows beyond the capsule. In one embodiment, the device comprises a semipermeable housing containing viable cells which produce and secrete the pre-selected molecule. The viable cells may be disposed on a flat surface such as a disk. In another embodiment, the cells are contained within porous tubes that allow the physiologic activity of the cells to occur and from which the preselected molecule or cell products can flow into the bloodstream or body fluids. The tubes may extend beyond the capsule.
Embodiments of the present invention will now be described in greater detail with reference to the attached drawings, which are provided for purposes of illustration and are not meant to be limiting of the scope of the invention. Referring to the drawings,
The viable cells may be attached to an inner surface of a fiber. Whether the cells are attached to the inner surface will depend upon the cell type included in the device. For example, some cell types grow in an anchorage dependent manner upon a solid surface while other cell types have no anchorage dependency and grow in suspension. The choice of cell type, however, is dependent upon the desired application.
The device of
A detail view of the disk 38 is illustrated in
In another embodiment of the invention, illustrated in
The body includes a manifold (illustrated in
As illustrated in
Alternatively, the end of the tube is not blocked and the filtrate is allowed to bath the cells and has a sufficient flow rate such that the cells are provided with the proper amount of nutrients and the unfiltered blood is not allowed to travel upstream to the cells.
As mentioned above, anchoring systems that may be used in the instant invention include devices that provide blood clot filtering or blood vessel stents. Useful anchoring elements are characterized by their ability to be anchored within the lumen of a conduit without occluding or preventing blood flow. One of the advantages of the present invention is that the extracorporeal device anchoring system does not have to be designed to be secured in a blood vessel without damaging the blood vessel. This minimizes the concern for tissue damage.
Blood clot filters are used routinely by medical practitioners to prevent the migration of potentially life threatening blood clots within the vasculature. Blood clot filters typically are designed to be implanted and anchored within the lumen of a blood vessel. When implanted, the anti-migration filters permit blood in the vessel to pass while simultaneously trapping blood clots. The devices and techniques useful toward attaching a filter to a blood vessel can be applied to extracorporeal devices. Of course, because the device is not intended to be introduced into the vasculature, the design of the anchors need not have the high level of precision required for their use.
Filters described in U.S. Pat. Nos. 4,817,600 and 5,059,205, referred to in the art as Greenfield® filters and available from Medi-Tech®, Boston Scientific Corporation, Natick, Mass., are particularly well suited to the practice of the invention. The cone-shaped Greenfield® vena cava filters are designed to provide maximal entrapment area for trapping blood clots while maintaining patency of the blood vessel after trapping emboli. The spacing between the six legs of the Greenfield® vena cava filters ensures the trapping of emboli greater than 3 mm (Greenfield et al. (1989) “Venous Interruption” Chapter 68, pp. 929-939 in “Haimovici's Vascular Surgery Principles and Techniques, Third Edition,” Appleton and Lange, Norwalk, Conn./San Mateos, Calif.). Accordingly, the filters may be able to capture capsules greater than 3 mm in diameter. In another embodiment of the invention, the device is incorporated into a tube that is inserted into an extracorporeal conduit. In this embodiment, the anchors are unnecessary because the device is secured within the tube.
The extracorporeal drug delivery device of the various embodiments of the present invention may be capable of delivering a preselected drug or cell products over an extended period of time. Because the device is not designed to be implanted, the size of the capsule can be scaled for higher delivery rates by increasing the size of the capsule and the drug delivery media.
Because the device permits delivery of the preselected molecule over defined periods of time, another important consideration in the design of the device is the configuration that will maintain the viability of the cells enclosed in the device. It is understood that a variety of factors, for example: the supply of oxygen and nutrients to the cells in the capsule; the removal of waste products from the cells in the capsule; the minimization of host immune responses directed against the cells in the capsule; the proliferative activity of the cells; and whether cells located at the center of the capsules are susceptible to pressure necrosis, all of which may influence the design and preparation of a cell containing tube(s).
In addition to adequate aeration, it is important that the encapsulated cells obtain sufficient amounts of essential nutrients from the blood supply to remain viable. The transport oxygen is almost entirely by diffusion and this is a significant aspect in maintaining cell viability. Once the geometry of a hollow fiber has been optimized for oxygen transport, then the hollow fiber inherently will be able to permit the diffusion of adequate amounts of nutrients into the lumen of the capsule from the blood stream. Similarly, such a geometry is contemplated also to permit diffusion of cell metabolites, including, waste products and the preselected molecule, out of the hollow fiber and into the blood stream.
The hollow fibers preferably are produced from a semi-permeable membrane having pores dimensioned to permit the diffusion of oxygen and nutrients into the lumen of the hollow fiber while permitting the efflux of cellular waste products and the pre-selected molecule out of the hollow fiber. In addition, the pores preferably are dimensioned to exclude the passage of cells therethrough. Accordingly, the pores are designed to prevent migration of the viable cells from the lumen of the hollow fiber into the blood steam, thereby maintaining the implanted cells at a single location in the host to facilitate their subsequent removal if or when necessary. The pores also are designed to prevent the influx of the hosts immune cells, for example, macrophages and lymphocytes, which if allowed to enter the lumen of the hollow fibers may be detrimental to the viability of the cells enclosed therein. The membrane, therefore, provides an immuno-protected environment that protects cells enclosed therein from an immune response. This may be an important consideration if the implanted cells are non-autologous in nature. If autologous cells are used per size restrictions longer than molecular dimension then the design would be modified accordingly.
The hollow fibers comprising, or for incorporation within, the capsule may be produced from biocompatible polymers which include, but are not limited to, polyvinylchloride, polyvinylidene fluoride, polyurethane isocyanate, polyalginate, cellulose acetate, cellulose diacetate, cellulose triacetate, cellulose nitrate, polysulfone, polystyrene, polyurethane, polyvinyl alcohol, polyacrylonitrile, polyamide, polymethylmethacrylate, polyethylene oxide, polytetrafluoroethylene, or copolymers thereof. A summary of currently available hollow fibers, including methods of manufacture and the names of commercial suppliers, is set forth in Radovich (1995) “Dialysis Membranes: Structure and Predictions” Contrib Nephrol., Basel, Karger, 113:11-24, the entire disclosure of which is incorporated herein by reference. In addition, polytetrafluorethylene polymer hollow fibers are available commercially from Impra, Inc., Tempe, Ariz. or W.L. Gore and Associates, Flagstaff, Ariz. U.S. Pat. No. 6,716,208, the contents of which is incorporated by reference, provide more details of the components of the device.
The disks provide a suitable material onto which the cells can be disposed. In a preferred form, the disks have a trabecular structure that allows the cells to grow into the medium. Alternatively, the cells may be disposed on the surface of the membrane. The disk material may consist of a variety of different types of compounds, including ceramics, carbon, and metallic substances. In one embodiment, the disks are made from carbon material coated with niobium (Nb). In certain embodiments, the disk material may be a porous, metal-coated reticulated open cell foam of carbon-containing material, the metal coating selected from, but not limited to, tantalum, titanium, platinum (including other metals of the platinum group), Nb, hafnium, tungsten, and combinations thereof. The coatings (which may be non-metal as well) also may include, but are not limited to, silica, nickel, alumina, gold, collagen, non-siliceous materials, and nanofibrous scaffolds, complexes, composites, or constructs formed by any variety of methods including, but not limited to, nano-magnetic deposition, nanocasting, fiberbonding, electrospinning, or any variation on these techniques. The coatings can increase integrity and strength of the disks. Alternatively, other shapes of substrates on and/or in which cells are disposed may be used, such as rectangular or square substrates. Further, the substrates can be made from various polymers, which may have a trabeculated structure. The polymers can promote or increase cell expansion, attachment, and/or viability.
Cell therapy presents itself as a new approach to the treatment of acute and chronic diseases. This therapeutic approach has its origins in the growing appreciation that most disease processes are not due to the lack of a single protein but develop due to alterations in complex interactions of a variety of cell products. Cell therapy depends on cell and tissue culture methodologies to expand specific cells to replace important differentiated processes deranged or lost in various disease states. Recent approaches have made progress by placing cells into hollow fiber bioreactors or encapsulating membranes as a means to deliver cell activities to a patient, requiring complex extracorporeal pump systems and large bioreactor devices. One clinical disorder that may be treatable with cell therapy is the systemic inflammatory response syndrome, or SIRS. It is a catastrophic sequela of a variety of clinical insults, including infection, pancreatitis, and cardiopulmonary bypass, and claims over forty thousand lives in the U.S. each year. The most common cause of SIRS is bacteria-induced septic shock. The exceptionally high mortality associated with the syndrome is due in part to the development of the highly lethal multiple system organ failure syndrome (MOF) in a subset of patients with SIRS.[references 1,2]
One area that presents opportunities is the development of miniaturized cell therapy devices containing renal tubule cells, which play an important immunologic regulatory role in septic shock. The successful development of a prototype is an important step toward allowing cell therapy to be delivered in a simple extracorporeal cassette without extracorporeal pump systems, thereby expanding this therapy to a broader indication, especially early severe sepsis syndrome prior to established acute renal failure (ARF).
An extracorporeal device utilizing a standard hemofiltration cartridge seeded with approximately 108 renal tubule cells grown from adult stem/progenitor cells as confluent monolayers along the inner surface of the fibers has been successfully fabricated with human cells.[reference 4] This initial cell therapy device is large (12×4-in cylinder) and requires an additional extracorporeal pump circuit to deliver blood and plasma ultrafiltrate to the cell-containing device. These elements were designed as an add-on to current dialysis treatment in patients with ARF.
The successful creation of the miniature device proposed has the potential to lead to a variety of other cell therapy devices including “wearable artificial organs”. Current cell therapy approaches target the use of stem cells for neurodegenerative disorders (Parkinson's, Alzheimer's), spinal cord injury, heart disease (congestive heart failure, myocardial infarct), pancreas disease (diabetes), liver disease (cirrhosis, hepatitis), kidney disease (end-stage renal disease [ESRD], ARF), blood (sickle cell anemia), muscle disorders (muscular dystrophy), skin (burns), and bones (arthritis, osteoporosis). From a market standpoint, the value of cell-based markets is estimated to be $26.5 billion in 2005, $56.0 billion in 2010, and $96.0 billion in 2015. More than 300 companies are involved in cell therapy. The near-term market for the use of renal cell devices is also very large and includes (U.S. only): ARF (100,000 patients, $2 billion market), ESRD (400,000 patients, $20 billion cost), and cardiorenal syndrome (350,000 patients, $25 billion cost).
Any of the embodiments of devices for use in extracorporeal circuits according to the invention (whether or not utilizing pumps and/or fluid separation devices) can be used to treat the disease states described herein (such as the cell therapy targets described above). Moreover, any type of cell can be used with any of the embodiments of devices for use in extracorporeal circuits according to the invention, including, for example, a eukaryotic cell, such as a mammalian cell, such as a renal cell (e.g., a renal tubule cell). The cells can treat bodily fluids in order to prevent or alleviate the disease conditions described herein.
The embodiment illustrated in
The materials of the embodiment illustrated in
The substrate in this embodiment can be shipped to the site of use with cells grown on the substrate and in a cryopreserved state. When the medical professional wishes to use the device, the substrate, and the cells thereon, can be thawed and loaded into the device. This configuration facilitates ease of manufacturing, shipping, and storage, as the pre-grown cells on the substrate can be shipped and stored in a frozen state.
Scaffolds 196 retain the disks within the interior space of the housing 184, and are shaped like a tuning fork in this embodiment. There are five scaffolds shown in this embodiment, but devices may have fewer or more of the them than shown. Additionally, each scaffold may hold any number of disks. A flow separator 202 within the interior of the housing 184 is disposed between the inlet 186 and the disks. When fluid enters the housing 184 through the inlet 186, the flow separator 202 provides desirable flow dynamics to allow the fluid to pass through the disks and adjacent the viable cells. Additionally, baffles 192 are disposed between the inlet 186 and the disks to further direct the fluids through the disks. Each of the baffles 192 is adjacent or proximate to a scaffold 196. The housing 184 also includes ports 194. The ports 194 allow for pressure release, fluid expansion, extracorporeal flow, and analysis probes. Additionally, the device may have fewer, more, or no ports. In this figure, the inlet 186, outlet 188, and ports 194 are luer fittings. These components can be selectively attached or detached from tubing or grooves. Alternatively, the inlet 186, outlet 188, and ports 194 need not be luer fittings. In certain situations, an adhesive (not shown) can be used with luer fittings, such as the inlet 186, outlet 188, and ports 194 in this figure, to further maintain a seal.
The embodiment illustrated in
The materials of the embodiment illustrated in
The embodiment of the present invention illustrated in
The embodiments of the present invention illustrated in
Devices of the present invention as illustrated in
The embodiments of the present invention illustrated in
Viable cells 300 are seeded within the ICS 1140 of at least one hollow fiber 1152. The cells 300 can be adherent cells. In certain embodiments, the cells 300 are grown on particles 298, such as beads or other microcarriers, and are seeded into the ICS 1140 while on the particles 298. Examples of microcarriers include, but are not limited to, HyQ® Spheres™ microcarriers (available from HyClone, Inc., South Logan, Utah and SoloHill Engineering, Inc., Ann Arbor, Mich.), Cultispher® microcarriers (available from HyClone, Inc. South Logan, Utah), Cytodex® microcarrier beads (available from GE Healthcare Bio-Sciences Corp., Piscataway, N.J.), and Glass microcarrier beads (available from Sigma-Aldrich, Inc., St. Louis, Mo.). The HyQ® Spheres™ microcarrier is a cross-linked, polystyrene microcarrier. It can be optionally coated with collagen or animal-derived component free materials. Further, it can have a cationic charge layering surface modification. The Cultispher® microcarriers have larger surface areas for attachment and are dissolvable for cell retention. They come in three types (G: basic gelatin cross linked, S: large pore cross linked, and GL: high temperature cross-linked). The Cytodex® microcarrier bead is a dextran, cross-linked microcarrier. It has positively charged Diethylaminoethyl (DEAE) groups throughout the matrix. The Glass microcarrier beads can be of varying size and density.
When the cells 300 are within the device 290, they are located within a biomatrix material 292, which itself is disposed within the ICS 1140. Thus, the cells are suspended in the biomatrix material 292 or physiologic buffer (e.g., a 3-D matrix suspension) in the hollow fibers 1152. Examples of a biomatrix material 292 or physiologic buffer include, but are not limited to, alginate solution, gelled alginate solution, nutragen, matrigel, collagen I or any three-dimensional forming or formed biocompatible matrix.
The hollow fibers 1152 of this or other hollow-fiber embodiments of the invention can have any of the following design characteristics. The hollow fibers can have immunoprotective properties, and cells can be cultured on the inside of the fibers. The pores through the hollow fibers can be of a size within the range of about 0.01 μm to about 0.1 μm. Pore sizes also can be measured as a molecular weight cutoff (MWCO) and can be of a type useful for ultrafiltration. MWCO and size are loosely related with 0.01 μm equating to a MWCO of about 70,000 kD and 0.1 μm equating to a MWCO of about 500,000 kD. The hollow fibers typically have a MWCO of about 40,000-50,000 kD. The hollow fibers can function as a barrier between the nutrient supply source and the cells.
The biomatrix material and/or cells can be inserted into the device at the point of manufacture or, alternatively, at the clinical site. If the cells are inserted at the point of manufacture, the device can be frozen and delivered in such a state. Then, the entire device is thawed prior to use by the medical professional. If the cells are inserted at the clinical site, the cells can be cryopreserved separately from the device. Then, the cryopreserved cells are thawed at the clinical site, when they are to be used, mixed with the biomatrix material and injected (e.g., with a syringe) into the device. Alternatively, the cryopreserved cells can be pre-mixed with the biomatrix material and both components can be thawed when used. In one embodiment, cryopreserved cells, grown on microcarriers, are thawed and, if necessary, mixed with a three-dimensional biomatrix, such as alginate. This mixture is injected into the device, and the mixture is allowed to gel. In the case of alginate, the addition of calcium chloride is beneficial for gel formation, although this step will vary depending on the biomatrix material and does not limit any three-dimensional matrix utilization. Alternatively, a three dimensional matrix may be excluded.
The materials of the embodiment illustrated in
The extracorporeal devices described herein can be used with extracorporeal circuits as illustrated in
The circuits of
The premise of this cell-based therapy is that as the ultrafiltrate passes through the cell chamber of the treatment device, systemic small-molecular-weight molecules in the ultrafiltrate, which have been produced due to a disease state, such as a septic disease state, trigger a therapeutic response in the cells within the treatment device, which in turn secrete small-molecular-weight molecules into the passing ultrafiltrate. This processed ultrafiltrate is then refiltered across the hollow fiber membrane of the second hemofilter and is then joined to the filtered blood and returned to the systemic circulation, with the released cell factors providing therapeutic value, for example, to the septic state.
In alternative extracorporeal circuits, such as those shown in
In
The circuit of
In an alternative extracorporeal circuit, shown, for example, in
Additionally, the lines of the circuit shown in
An alternative extracorporeal circuit is shown in
Practice of the invention will be still more fully understood from the following examples, which are presented herein for illustration only and should not be construed as limiting the invention in any way.
Early prototype formulation of miniaturized cell therapy devices are schematized in
Fabrication and in vitro testing of one prototype, shown in
Fabrication and in vitro testing of a second prototype with cryopreservation storage capacity. Prototype development of a second cell therapy device is also schematized in
The extracorporeal device of various embodiments of the present invention can be used with a blood or fluid conduit as illustrated in
It is understood, however, that the physicians judgment based upon clinical circumstances should determine on a case by case basis the optimal mode for introducing the device as well as the optimal location for anchoring the device. Such judgments are contemplated to be within the scope of expertise of the skilled physician.
Practice of the invention will be still more fully understood from the following examples, which are presented herein for illustration only and should not be construed as limiting the invention in any way.
Two different device designs are envisioned for the therapeutic delivery system:
An immortalized renal cell line (MDCK) was seeded unto Nb-coated carbon based disks. Prior to cell seeding, disks were coated with either collagen I, collagen IV, laminin, or matrigel. Seeding densities were 106 cells/disk. The disks were maintained in a spinner culture for 1 week. Cell-seeded disks were then cryopreserved, stored in liquid nitrogen and then thawed, with a non-stir rest period under tissue culture conditions of 2-4 hours post thaw. Disks were then reinstated in spinner culture bottles and assayed for cell density 24 hours post thaw. Lactate production was utilized to determine cell expansion on the disks. MDCK cell densities on disks with various bio-matrix coatings pre and post cryopreservation are depicted in the graph of
In order to ascertain if primary cells could also be maintained on the Nb-coated carbon based disks, primary porcine renal proximal tubule cells (PPTC) were seeded on disks coated with various bio-matrixes (either no matrix, collagen IV, laminin, or matrigel) at 106 and 107 cells/disk. Cell-seeded disks are currently being maintained in spinner culture. Primary porcine renal tubule cell growth rates on disks, as determined by lactate production, are depicted
Cells from an immortalized renal cell line (MDCK) resuspended in either 1) nutragen (collagen I (3.12 mg/mL) gel supplemented with laminin (100 μg/mL) and collagen IV (100 μg/mL)) or 2) matrigel were seeded into hollow fiber bioreactors at a density of 2.06×107 cells/mL of gel. Cell loaded bioreactors were maintained in culture for 8 (matrigel) and 13 (nutragen) days, with cell density assessed via lactate production. Results are depicted in the
This example describes the design of the device illustrated in
Tissue engineering and cell therapy is a new and exciting approach to the treatment of acute and chronic diseases. The potential success of this therapeutic approach lies in the growing appreciation that most disease processes are not due to the lack of a single protein but develop due to alterations in complex interactions of a variety of cell products. Cell therapy is dependent on cell and tissue culture methodologies to expand specific cells to replace important differentiated processes deranged or lost in various disease states. Recent approaches have made progress by placing cells into hollow fiber bioreactors or encapsulating membranes as a means to deliver cell activities to a patient. A reasonable extension of this approach is to add cell therapy to the current successful renal substitution processes of hemodialysis and hemofiltration.
The research outlined herein is a new therapeutic approach based upon the thesis that renal tubule cells play an important immunologic regulatory role in septic shock. Accordingly, the effect of renal tubule cell therapy improves cardiovascular parameters, immunologic activity, and ultimate survival rate of normal, non-uremic large animals in septic shock. The approach is based upon the observation that severe septic shock results in ATN and ARF within hours of bacteremia in a porcine model of septic shock. Thus, ARF develops early in the time course of septic shock, a time frame not appreciated clinically since it takes several days to observe a rise in blood urea nitrogen (BUN) and serum creatinine after the acute insult. Experiments suggest that septic shock induces renal tubule injury, loss of immunologic balance, and cardiovascular collapse with declines in solid organ perfusion. This proposal will develop a device prototype to deliver cell therapy in simpler extracorporeal perfusion systems with greater distribution and storage capabilities for on-site, on-demand availability.
The disorder of ARF, or ATN, may be especially amenable to cell therapy in conjunction with continuous hemofiltration techniques, since acute hemodialysis or hemofiltration has yet to reduce the mortality rate of ATN below 50 percent, despite advances in synthetic materials and extracorporeal circuits. ATN develops predominantly due to the injury and necrosis of renal proximal tubule cells. The early replacement of the functions of these cells during the episode of ATN, which develops concurrently with septic shock, will provide almost full renal replacement therapy in conjunction with hemofiltration. The addition of metabolic activity, such as ammoniagenesis and glutathione reclamation, endocrine activity, such as vitamin D3 activation, and cytokine homeostasis is contemplated to provide additional physiological replacement activities to change the current natural history of this disease process.
In this regard, the cause of death in patients with acute renal failure is usually the development of SIRS, often secondary to sepsis, with resulting cardiovascular collapse, ischemic damage to vital organs, and MOF. The propensity of patients with ARF to develop SIRS and sepsis suggests that renal function, specifically renal tubule cell function secondary to ATN, plays an important immunomodulatory role in individuals under stress states. The renal tubule cell's roles in glutathione reclamation, glutathione peroxidase synthesis, and activation of vitamin D3, with its important immunoregulatory functions, are pathways to maintain tissue integrity and host defense under stress conditions. A less recognized role of the kidney, and the renal tubule cells, is its potential immunoregulatory function. The kidney is derived embryologically from dorsal mesoderm, a collection of cells also important in the development of bone marrow stem cells. The maturation of cells responsible for erythropoietin synthesis and activation of 1,25-(OH)2 vitamin D3 in the kidney is reflective of this embryonic origin. Phylogenetically, in bony fish and amphibians without lymph systems, the kidney is the major antibody-producing organ. Not surprisingly, mammalian renal proximal tubule cells are immunologically active. They are antigen-presenting cells that possess co-stimulatory molecules and that synthesize and process a variety of inflammatory cytokines.
ARF, and ATN that results in the loss of the kidney's immunoregulatory function, results in a propensity to develop SIRS, sepsis, MOF, and a high risk of death. A rise in sepsis events can be observed from 3% to nearly 60% in patients who develop ARF during the post-op course following open heart surgery. A counterpart to this loss of immunologic function has been observed in chronic renal insufficiency and end-stage renal disease, which are clearly pro-inflammatory states. The degree of inflammation in these patient populations has been highly correlated to mortality rates. The loss of renal tubule cells, rather than loss of filtration and clearance function, may be the cause of this inflammatory dysregulation observed in these patients as well.
In the past decade, a large constellation of data has provided new insights into the inflammatory response that seems to underlie the MOF syndrome. There are now data linking patient outcome to initial plasma levels of tumor necrosis factor (TNF)-a, interleukin (IL)-6, and other proinflammatory cytokines. There is also increasing agreement that organ dysfunction in sepsis arises from oscillating patterns of inflammation from systemic spread of mediators beyond their usual autocrine or paracrine pathways, alternating with immunosuppression either as compensation for or exhaustion of the inflammatory response. This futile and lethal state of affairs has been referred to as “immunologic dissonance.” This paradigm has led to efforts to interrupt this process by targeting individual elements in the cascade or by reducing the overall burden of inflammatory mediators by plasma replacement or adsorption. Other approaches, including plasmapheresis, plasma exchange, hemofiltration, or monoclonal antibodies directed at various components of the inflammatory cascade, have not consistently demonstrated clinically useful interruption of the inflammatory cascade, despite promising in vitro and animal studies.
Bioartificial Renal Tubule Assist Device (RAD)
An extracorporeal device was recently developed that utilized a standard hemofiltration cartridge seeded with approximately 108 renal tubule cells which grow as confluent monolayers along the inner surface of the fibers. The nonbiodegradability and the pore size of the hollow fibers allow the membranes to act as both scaffolds for the cells and as an immunoprotective barrier. In vitro studies of this renal tubule assist device (RAD) have demonstrated that the cells retain differentiated active transport properties, differentiated metabolic activities, and important endocrine processes. See Humes et al., Metabolic Replacement of Kidney Function in Uremic Animals with a Bioartificial Kidney Containing Human Cells, Am. J. Kidney Dis. 39(5):1078-1087 (2002). Additional studies have shown that the RAD, when incorporated in series with a hemofiltration cartridge in an extracorporeal blood perfusion circuit, replaces filtration, transport, metabolic, and endocrine functions of the kidney in acutely uremic dogs. Id. Furthermore, other recent studies have demonstrated that the RAD ameliorates endotoxin shock in acutely uremic animals. RADs consisting of both porcine and human cells have been successfully fabricated and tested.
The bioartificial kidney set-up consists of a filtration device (a conventional hemofilter) followed in series by the tubule RAD unit. Specifically, blood is pumped out of a large animal using a peristaltic pump. The blood then enters the fibers of a conventional hemofilter, where ultrafiltrate is formed and delivered into the lumen of the hollow fibers within the RAD, downstream to the conventional hemofilter. Processed ultrafiltrate exiting the RAD is collected and discarded as urine. The filtered blood exiting the hemofilter enters the RAD through the ECS port and disperses among the fibers of the device. Upon exiting the RAD, the processed blood travels through a third pump and is delivered back to the animal. This additional pump is required to maintain appropriate hydraulic pressures within the RAD. In this regard, the pressures of the blood and ultrafiltrate just before entry into the RAD are monitored. Heparin is delivered continuously into the blood before entering the RAD to diminish clotting within the device. The RAD is oriented horizontally and placed into a temperature-controlled environment. The temperature of the cell compartment of the RAD must be maintained at 37° C. throughout its operation to ensure optimal functionality of the cells. Maintenance of a physiologic temperature is a critical factor in the functionality of the RAD. The unit is able to maintain viability because metabolic substrates and low-molecular-weight growth factors are delivered to the tubule cells from the ultrafiltration fluid and the blood in the ECS. Furthermore, immunoprotection of the cells grown within the hollow fiber is achieved due to the impenetrance of immunoglobulins and immunologically competent cells through the hollow fibers. Rejection of the cells, therefore, does not occur. This arrangement thereby allows the filtrate to enter the internal compartments of the hollow fiber network, lined with confluent monolayers of renal tubule cells for regulated transport and metabolic function.
Improvements to Renal Cell Therapy Devices
The RAD is an extracorporeal device utilizing a standard hemofiltration cartridge containing renal tubule cells grown from adult stem/progenitor cells as confluent monolayers along the inner surface of the fibers. This initial cell therapy device requires large cell numbers for seeding, is labor intensive to generate and maintain and cannot be held in a frozen state for ready deployment at a clinical site. In addition, the RAD is large (12×4-in cylinder) and requires an additional extracorporeal pump circuit to deliver blood and plasma ultrafiltrate to the cell-containing device. The RAD was designed as an add-on to current dialysis treatment in patients with ARF. The clinical use of renal tubule cell therapy for patients with severe sepsis may not require this complex two-cartridge system with two extracorporeal pump systems, since most of these patients are not in ARF. Accordingly, a miniaturized cell therapy device will not require extensive extracorporeal blood pump systems. In addition, a miniaturized device that could be stored at the clinical site for immediate use is desirable to succeed as a commercial product. The current RAD is stored at a central manufacturing facility at 37° C. and must be shipped at 37° C. to the clinical site, delaying treatment and adding to the cost of therapy. Development of a cell device that can be cryopreserved and/or stored at clinical sites is a desirable feature for optimal deployment of such devices. The research described herein will develop a prototype for these more flexible device requirements for easier clinical implementation. As detailed below, this research will develop 1) a miniaturized bioartificial renal cell system (BRECS-d) device that can be manufactured under conditions that allow for cryopreserved cell stocks to be stored at clinical sites for easy deployment and 2) a simplified, bloodless circuit for administration of BRECS-d therapy. These improvements will enable use of the technology for a broader number of indications. Results of the work done with the larger RAD system served as initial proof-of-principle for these miniature BRECS-d devices.
A 10-16 Hour Porcine Gram-negative Model of Septic Shock
Based on severe septic shock results in ATN and ARF within hours of bacteremia in a porcine model of septic shock (as indicated by cessation of urine output and renal histology), the effect of renal tubule cell therapy improves cardiovascular parameters, immunologic activity, and survival rate of normal, non-uremic large animals in septic shock. Thus, ARF develops early in the time course of septic shock, a time frame not appreciated clinically since it takes several days to observe a rise in BUN and serum creatinine after the acute insult. Septic shock induces early renal tubule injury, loss of immunologic balance and cardiovascular collapse with declines in solid organ perfusion.
RAD therapy at the onset of bacteremia (RAD therapy is initiated at the same time that bacteria is infused) in these established animal models “prevents” the rapid deterioration from SIRS. To assess more effectively the efficacy of renal cell therapy to treat sepsis and septic shock, a more reproducible porcine model of septic shock, in which the time to death is greater than 4-6 hours after initiation of sepsis, needs to be developed.
Results
A reproducible model of SIRS in a porcine model of sepsis to test various innovative approaches to treat this clinical disorder has been established and published. See Humes et al., Cell Therapy With a Tissue Engineered Kidney Protects Against the Multi-Organ Consequences of Septic Shock, Critical Care Med. Vol. 31, at 2421-2428 (2003). To improve this animal model, the model was refined to be more consistent and reproducible. The model was also extended to the tissue-specific effects of SIRS after potential therapeutic interventions were attempted. The published model had a wide variability in cardiovascular parameters and time to death (3-7 hours in sham control and 7-12 hours in cell-RAD treatment groups). To improve the reproducibility of this model, volume resuscitation protocol was increased from 100 mL/hr to 150 mL/hr immediately after the crystalloid/colloid bolus infusion at the time of bacteria administration, as detailed below. In addition to this improved resuscitation protocol, the tissue-specific consequences of sepsis with or without RAD therapy were evaluated with bronchoalveolar lavage (BAL) to better understand the immunoregulatory role of renal tubule cell therapy. BAL specimens were used to assess pulmonary microvascular damage and inflammation in response to SIRS. Data detailed below demonstrated that renal cell therapy was associated with less protein leak from damaged blood vessels and less inflammation in BAL fluid samples, as well as improvement of other cardiovascular effects of SIRS.
Pre-Clinical Large-Animal Model
The improved porcine model of septic shock was used to preliminarily assess the multiorgan effects of RAD intervention. The following data are derived from a total of 14 animals: 7 sham controls and 7 RADs (second cartridge containing porcine renal epithelial cells).
Cardiovascular Parameters
As demonstrated in
These data with this improved porcine model confirm that the RAD improved cardiac performance, renal blood flow (data not shown), and survival times compared to sham controls.
To investigate the effect of various therapeutic interventions in this model of SIRS, BAL fluid was obtained at the time of death and evaluated for protein content as a parameter of microvascular damage, various inflammatory cytokines and the absolute number of polymorphonuclear cells (PMNs). As summarized in Table 1 below, data suggested that RAD treatment resulted in less vascular damage and protein leak and less inflammatory cytokine release in the early phase of pulmonary involvement in SIRS. Levels of IL-6, IL-8 and TNF-α were lower in the treatment interventions versus sham controls. Levels of IL-1 and IL-10 (not shown due to no detectable levels in each group) were not different. Absolute neutrophil counts in the sham controls were above 1000 cells/mL, and the RAD group trended lower, although the n=2 in each group does not allow for statistical analysis.
BAL sampling was initiated for a subset of animals to assess the utility of this analysis. With these encouraging results, the next steps will incorporate this as a routine inflammatory index at 3-4 hours post-bacteria infusion. Data for plasma levels of various cytokines also showed substantive effects in the RAD treatment group compared to sham controls. At 6 to 8 hours post-sepsis induction, plasma levels for IL-10 were 59±13 and 49±16 (p=0.03); IL-8, 602±550 and 168±165 (p=0.01); and IL-6, 8.3±3.4 and 9.0±2.6 (p>0.05) for the RAD (n=7) and sham controls (n=7), respectively. The higher levels of IL-8 may reflect the effect of sepsis in increasing IL-8 release from renal epithelial cells.
In summary, a more reproducible porcine model of sepsis was developed with longer survival times to compare therapeutic efficacy of new device formulations, which will be the focus of the next research steps. In addition, tissue-specific inflammatory effects of therapy with BAL fluid analysis were incorporated into this model system.
This research will develop further preclinical data to demonstrate that cell therapy with a RAD can also reverse and improve cardiovascular performance, cytokine balance and survival times in the porcine model of E. Coli-induced septic shock discussed above. As opposed to the established “prevention” model of septic shock described above, which introduces sepsis and RAD therapy simultaneously, this proposed model will more closely simulate the clinical setting. Septic shock will be induced 2-4 hours prior to initiation of RAD therapy. This model will assess the ability of the RAD to reverse the process of septic shock, as measured by increased survival time, hemodynamic stability and reestablishment of cytokine balance.
Results
With the development of this improved porcine model, cell therapy interventions were introduced 2-4 hours after induction of sepsis to evaluate the efficacy of the RAD in reversing the progression of septic shock. Several animals were tested with the introduction of the RAD into the extracorporeal circuit 2-4 hours into sepsis. The introduction of the RAD and its accompanying blood lines resulted in such a drastic after-load reduction maneuver that all animals died within one hour of device placement, despite large volume resuscitation efforts with both acellular and cellular devices. Thus, this acute model system was not adequate to assess the potential for reversibility of septic shock progression. The results of ongoing clinical trials with human cell RADs, however, have presented data supportive of reversing the progression of sepsis in patients with ARF and MOF.
Current extracorporeal treatment with the RAD requires blood to be circulated into the ECS of the RAD while ultrafiltrate (UF) from the synthetic hemofilter is delivered to the luminal spaces in which the cells reside in the RAD, duplicating the structural and functional anatomy of the nephron. Blood perfusion in the ECS makes the treatment with the RAD more complex and, perhaps, less safe (due to clotting and increased patient blood volume required by RAD circuit). Utilizing UF delivery alone to the RAD may provide sufficient pathophysiologic signals to the renal tubule cells to exert their immunoregulating effects during SIRS. A series of experiments were established to evaluate whether the biochemical effects of RAD therapy and potential amelioration of the natural history of SIRS in this porcine model can be accomplished with removal of blood perfusion within the RAD circuit.
Results
Work focused on the in vitro development, followed by a preliminary proof-of-concept pre-clinical animal testing, of a miniaturized BRECS to replace renal metabolic and hormonal processes. The successful employment of the BRECS unit will enable use of this technology for a broader number of indications, including utilization of a blood-free delivery system. Results of the work done above with the larger renal assist device (RAD) serve as initial proof-of-principle for the miniature BRECS unit. A BRECS unit that could be stored at the clinical site for immediate use will improve its ability to be used in a clinical setting. Development of a cell device that can be cryopreserved and stored at clinical sites is desirable clinically. The most advanced design of these BRECS prototypes (the carbon disk-based BRECS-d) was then tested in 2 animals, under the established porcine model of sepsis, utilizing an extracorporeal system that incorporated ultrafiltrate delivery, rather than blood, to the BRECS-d unit. The data below detail 1) the in vitro results that show the BREC-d design to allow for a robust cell system that can be cryopreserved and thawed with no detrimental effects on cell viability and 2) the incorporation of the BRECS-d unit in the septic animal model utilizing ultrafiltrate delivery, rather than blood, showing an improvement in survival time over both sham- and RAD-treated animals.
PPRECs were seeded on Nb-coated carbon-based disks at a density of 106 cells/disk. Prior to cell seeding, disks were coated with collagen IV. Preliminary studies determined that collagen IV—coated disks displayed superior cell growth as compared to disks coated with collagen I, laminin, or matrigel. The disks were maintained in a spinner culture for up to 7 weeks. After the initial post-seeding expansion, cell number, as determined by lactate production, remained stable over the 7-week period (
Cell growth coverage on the disks was assessed by both staining with crystal violet and DAPI.
To determine the feasibility of cryopreservation of the cell-seeded disks, the following study was performed. PPRECs were seeded on Nb-coated carbon-based disks, pre-coated with collagen IV, at a cell density of 106 cells/disk. The disks were maintained in a Starwheel spinner culture for two weeks. The Starwheel system was designed so that four disks were held by a tuning fork (i.e., a scaffold) in a vertical position with up to eight scaffolds per Starwheel flask. After two weeks, five scaffolds, each containing four disks, were placed in a BRECS-d unit for pre-cryopreservation oxygen and glucose consumption and lactate production rate determination. The cell-seeded disks were then cryopreserved in groups consisting of four disks per scaffold system, stored in liquid nitrogen for 48-72 hours, and then thawed for post-cryopreservation oxygen and glucose consumption rate assessment. The 4-disk groups were quick thawed at 37° C., and the five scaffold systems were placed in the BRECS-d unit illustrated in
Apart from these results, further in vitro testing of the BRECS-d unit, illustrated in
Upon the above promising in vitro studies shown in
The study results obtained from these two animals treated with the device shown in
The effect of PPRECs on normal porcine leukocyte activation was assessed. These studies isolated leukocytes from heparinized blood utilizing polymorph solution (per manufacturer's recommendation), followed by a 30-minute incubation period with or without LPS. During this 30-minute period, the leukocytes were incubated in fresh media or PPREC-conditioned media (exposed to PPRECs for 24 hours prior to removal for leukocyte incubation). Intracellular oxidative capacity was measured at the end of the incubation period by addition of 5 μM 2′,7′-dichlorofluorescein (DCFH)-diacetate. The PPREC-conditioned media significantly decreased leukocyte activation, as determined by oxidative burst, under both non-LPS and LPS conditions compared to the fresh media condition. The results are depicted in
Although the above-described leukocyte oxidative burst assay generated preliminary data that could be used to show the potential for assessing differences in the activated state of systemic leukocytes, this assay can be cumbersome and dependent on the isolation methods employed to obtain the leukocytes to be used in this assay. Variability was found due to the isolation process itself activating the leukocytes. An alternative method which could be used as an indicator of leukocyte activation has now been developed. The following protocol will be used for assessing the level of neutrophil, monocyte, and eosinophil activation in porcine peripheral blood. This approach measures the level of CD11b expression on the cell surface of each leukocyte population as an indicator of the “activation” level attained by the population. Neutrophils and to a lesser extent monocytes, and to an even lesser extent eosinophils, mobilize intracellular stores of CD11b to the cell surface as they become activated. The protocol calls for measuring the level of leukocyte surface expression by staining freshly isolated whole blood with a FITC-conjugated anti-porcine CD11b antibody. Red cells are then lysed and the leukocytes fixed using Becton-Dickenson's “Facs lyse” solution. This serves a three-fold purpose: it fixes the antibody-stained leukocytes for delayed flow cytometry analysis, it destroys the red blood cells, eliminating them from the analysis, and it separates the two granulocyte populations (neutrophils and eosinophils) from each other on the flow cytometer's forward and side scatter plots. Cells are then collected by centrifugation and resuspended for flow cytometry analysis. The data depicting the ability to assess the level of activation in various leukocyte populations is shown in
This assay will be used in two ways. First, the effect of cell therapy compared to sham controls on leukocyte activation will be assessed. Blood will be collected and leukocytes isolated at baseline (prior to bacteria infusion) and then again at 6 hours post-bacteria infusion. The intensity of fluorescent signal of FITC-conjugated anti-porcine CD11b antibody bound to the leukocyte surface will be analyzed by flow as an indicator of leukocyte activation and compared between sham- and cell-treated groups. The effect of an HREC-seeded BRECS-d device on leukocyte activation in an ovine anephric model is described in Example 7. Second, conditioned media will be collected from the BRECS-d units and incubated with normal, non-septic isolated porcine leukocytes under non-LPS and LPS exposure conditions. Anti-porcine CD11b antibody staining of the leukocytes will be assessed and used as a biomarker index of the BRECS unit's ability to inhibit the activation of leukocytes. It is contemplated that the BRECS unit will function similar to the kidney, by secreting various cytokines/molecules to mediate the activation of the leukocytes.
After encouraging pre-clinical animal data, including the use of the porcine SIRS model described previously, the FDA approved the initiation of a Phase I/II clinical trial to evaluate the safety and efficacy of this renal cell therapy utilizing the RAD on 10 critically ill patients with ARF and multiorgan failure receiving CVVH. See Humes et al., Initial Clinical Results of the Bioartificial Kidney Containing Human Cells in ICU Patients with Acute Renal Failure, Kidney Int'l 66(4):1578-1588 (2004). The predicted hospital mortality rates for these patients averaged greater than 85%. The devices used in this study were seeded with human renal proximal tubule cells isolated from kidneys donated for cadaveric transplantation but found to be unsuitable for transplantation due to anatomic or fibrotic defects. The results of this clinical trial demonstrated that the experimental treatment could be delivered safely under study protocol guidelines for up to 24 hours when used in conjunction with CVVH. The clinical data indicated that the RAD exhibited and maintained viability, durability, and functionality in this clinical setting. Cardiovascular stability of the patients was maintained, and increased native kidney function, as determined by elevated urine outputs, temporally correlated with RAD treatment. The device also demonstrated differentiated metabolic and endocrinologic activity. All but one treated patient with more than a 3-day follow-up showed improvement, as assessed by acute physiologic scores. Six of the 10 treated patients survived past 28-days with kidney function recovery, although mortality rates predicted for these 10 patients using the acute physiology, age, chronic health evaluation (APACHE) 3 scoring system were on average 85 percent. Plasma cytokine levels suggests that RAD therapy produces dynamic and individualized responses in patients depending on their unique pathophysiologic conditions.
These favorable Phase I/II trial results led to a subsequent FDA-approved, randomized, controlled, Phase II investigation at 12 clinical sites to determine whether this cell therapy approach alters patient mortality. This Phase II study involved 58 patients, of whom 40 were randomized to RAD therapy and 18 made up a control group with comparable demographics and severity of illness by SOFA scores. Of importance, the incidence of sepsis in this clinical trial was 73% in the RAD treated group and 67% in the control group. The efficacy of the RAD in patients with sepsis demonstrates the ability of renal cell therapy to reverse sepsis in the clinical setting. The early results were as compelling as the Phase I/II results. Renal cell therapy improved the 28-day mortality rate from 61% in the conventional hemofiltration-treated control group to 34% in the RAD-treated group. This survival impact continued through the 90- and 180-day follow-up periods (p<0.04), with the Cox proportional hazard ratio indicating that the risk of death was 50% of that observed in the continuous venovenous hemofiltration (CVVH) therapy group (
As shown above, the BRECS-d design allows for the device to be both miniaturized and cryopreserved, thus eliminating the shortcomings of the original RAD design, such as needing to maintain the RAD in culture prior to clinical use, which demanded significant hands-in technical time. The technology is a bioengineered device that utilizes renal cells in the treatment of diseases, such as SIRS and septic shock. The target treatment population includes patients with SIRS and septic shock. Based on the in vitro testing results shown in
As described above, the BRECS-d device of
The goal is to develop a prototype design to allow the entire BRECS-d unit to be cryopreserved. Additionally, the BRECS-d system will continue to be tested in vitro to assess cryopreserved shelf life, establish protocols for manufacturing, transportation and clinical deployment. The prototype development will determine the robustness of the unit design within the perceived use/lifecycle environment. The lifecycle process consists of the following: 1) injection molding of the unit, followed by sterilization and packaging, 2) sterile loading of the cell disk tuning fork cassettes into the unit, 3) sterile filling of the unit with freezing medium, and 4) sealing the cell holding unit for storage. The unit will then be stored in a sterile cryo-compatible bag filled with the same freezing medium. Once sealed, pressure ports will be opened to allow for fluid expansion, and the unit will begin a temperature step-down process, in reasonable degree increments that allow for immersion in liquid nitrogen for maintenance of maximum cell viability over an indefinite length of storage time. At point of use, the bagged unit is removed from its liquid nitrogen storage and placed into a bath of 37° C. water for thawing. Once the bagged unit is temperature stabilized, the pressure ports will be closed to maintain sterility as the unit is removed from the bag, dried, aseptically wiped, and connected to the ancillary circuit for patient use.
The resins being contemplated for injection molding are approved for biomedical use and have data showing compatibility with EtO and gamma radiation sterilization. Additionally, these materials have been used in cryo-storage tubes, but data have not been found showing performance at extended extreme low temperatures under the exact conditions of the envisioned prototype. The issues of thermal shock, stress cracking, and impact strength will be monitored, as well as differences between the coefficient of thermal expansion of the threaded fasteners and gasket material. The gaskets and fasteners are rated for the sterilization and low temperatures. The development process will address these parameters and provide data showing the robustness of the molded device's performance within its proposed environment.
Testing on the device will include the following.
Therefore, a total of 96 BRECS-d units (26+30+40) will be tested in the in vitro study plan. The above studies will use 8 PPREC isolations in order to have primary cells from multiple animals.
Methods
The Starwheel carbon-based disk cultures will be established utilizing the following protocol: 8×107 PPRECs per 10 mL media will be perfused under recirculating conditions into a flow through chamber containing collagen IV—coated disks (0.5 μg/cm2) to allow for the PPRECs to optimally infiltrate the center of the carbon disks. Disks are incubated at 37° C. for 3-4 h before being placed in the Starwheel Dynamatrix cell growth system. Cells are spun at 2 rpm and in 175 mL of media. Media is changed 3 times weekly by aspirating old media through side arm and adding new media. Side arm caps are loosened a quarter turn to allow for gas exchange in incubator. Disks are typically used after 2-5 weeks of culture. For initiation of the BRECS-d culture, Starwheel disk tuning forks, containing 4 cell-seeded carbon disks each, are sterilely transferred to the BRECS-d system. The BRECS-d unit will then either be connected to the in vitro perfusion system for culture and in vitro studies or prepared for cryopreservation. Tuning fork cryopreservation protocol will be followed for the initial BRECS-d cryopreservation protocols. The BRECS-d chamber will be filled sterilely with a freezing medium; care will be taken so that the fill level covers all disks but still allows for expansion during the freezing process. The unit will be placed in a sterile bag designed for cryopreservation and the bag will then be filled with freezing medium and sealed such that the BRECS-d unit ports can be manipulated and caps opened for release of any pressure build-up that could develop during the cryopreservation process. After opening of the ports, the BRECS-d cryo-system will then be reduced in temperature in a step-down process until transfer to liquid nitrogen. At the time of use, the BRECS-d cryo-system will be removed from liquid nitrogen storage and placed in a 37° C. waterbath for rapid thaw. The unit will then have the ports closed while still in the cryo-bag and will then be removed once port closure and device integrity have been visually confirmed. The BRECS-d will be sterilely connected to a perfusion circuit, utilizing techniques similar to those used for peritoneal dialysis catheter connection. The perfusion rate will be slowly increased until a desirable rate to sustain the nutritional and oxygen needs of the cells is achieved. Lactate will be measured using a commercially available kit (BioVision, Inc., Mountain View, Calif.). Glucose levels will be measured via a glucometer. Oxygen consumption will be measured online with, for example, a multi-frequency phase fluorometer that utilizes a fiber optic reflection oxygen probe and a ruthenium based or porphyrin complex sensor that works by fluorescence quenching of the complex to measure the partial pressure of dissolved or gaseous oxygen. Cell release will be assessed via manual hemocytometer cell counts and leukocyte activation via the DCFH assay described above.
Contemplated Results
A BRECS-d prototype will be designed and fabricated that can be cryopreserved and thawed. This design will allow for a BRECS-d unit that has therapeutic value. A shelf life of up to 1 year will be determined using established techniques for the cryopreservation of PPREC. These cells have been thawed after several years of cryopreservation with no noticeable negative effect.
A series of studies will be carried out to test the efficacy of a BRECS-d unit (post-cryopreservation) seeded with primary porcine renal epithelial cells (PPRECs) in the treatment of sepsis using a large-animal model of septic shock. A variety of hemodynamic parameters, systemic cytokines, BAL parameters, systemic myeloperoxidase (MPO), survival time, urine and ultrafiltrate NGAL levels (as an indicator of acute kidney injury (AKI)), systemic free-elastase levels, and leukocyte activation parameters will be assessed and compared in sham- and cell-treated groups.
Rationale
The following parameters will be measured in the large animal studies.
Research Plan
Two groups of animals will be evaluated and compared: Group 1 receiving BRECS-d therapy and Group 2 receiving a sham BRECS-d unit containing no cells. Each group will have 11 animals. Outcome measurements will include the above parameters in addition to survival time. Outcome will be considered positive if the BRECS-d units are shown to be effective in treating sepsis as indicated by an overall decreased inflammatory state (as indicated by systemic, BAL macrophage and circulating monocyte cytokine panels and MPO activity), increased cardiovascular and pulmonary function, increased survival time, and a decrease in leukocyte activation. An additional outcome measurement is to compare these above detailed parameters in BRECS-d units that have gone through the cryopreservation process to BRECS-d units that have not been cryopreserved.
Methods
Animal Study Design
The planned experiments this proposal are based on a well-established porcine model of bacteremic septic shock, which was further refined and in which death occurs in 8 hours or less. In brief, pigs weighing 30-35 kg will be used in these studies. All pigs will be treated for 7 days prior to therapy with the antibiotic Baytril in an attempt to eliminate any pre-existing infections that could skew experimental results. Baseline blood cultures will be drawn to assess the animal infection state prior to initiation of the study. Animals with positive baseline cultures that could affect study outcome will be eliminated from the study group set. After administration of anesthesia and intubation, pigs will undergo placement of dialysis catheter, arterial catheter, and Swan-Ganz thermodilution catheter (which are connected to transducers) to monitor arterial blood pressure, cardiac output, and central venous pressures. A Foley catheter is placed in the urinary bladder to measure hourly urine output. Doppler flow probes are placed around a renal artery and connected to a flow meter (available from Transonics Systems Inc.®, Ithaca, N.Y.). After all access lines are in place and secured and baseline parameters and laboratory studies obtained, 200 cc of broth containing 30×1010 bacteria/kg body weight of E. Coli (serotype 06:K2:H1) is instilled into the animals' peritoneal cavities. To better replicate the human clinical situation, cefriaxione (100 mg/kg) broad spectrum antibiotic will be administered 15 min after bacteria instillation. At this time, 80 mL/kg of crystalloid and 20 mL/kg of colloid (Hetastarch) is administered. Systemic heparinization is administered to maintain patency of the extracorporeal circuit. At the time of bacteria administration, conventional CVVH is initiated with a Fresenius F-40 hollow-fiber dialysis cartridge (HF-1; placed pre-BRECS-d) along with a BRECS-d unit and second F-40 cartridge (HF-2; placed post-BRECS-d), as depicted in both
Cardiovascular parameter measurements and samples for biochemical analysis are obtained at various time intervals during the experiment. At baseline and 6 hours post-bacteria infusion, a BAL will be performed for assessment of pulmonary microvascular damage and inflammation in response to SIRS. Leukocytes will be isolated at baseline and hourly post-bacteria infusion for assessment of the effect of cell therapy compared to sham controls on leukocyte activation. Urine and ultrafiltrate samples will be collected hourly for assessment of NGAL levels as an early indicator of acute kidney injury, and serum elastase will be measured to assess neutrophil degranulation.
Test Parameter Methods
Serum chemistries will be measured with an automated chemical analyzer. Cytokine levels will be measured with commercial ELISA assay kits reactive to porcine cytokines: IL-1β, IL-6, IL-8, IL-10, TNF-α, and IFN-γ (R&D Systems). BAL fluid will be obtained for cell counts and cell-type distribution, protein as a measure of vascular leak, myeloperoxidase (MPO) as an indicator of the activated state of neutrophils that have infiltrated the lung air space from the lung tissue, elastase as indicator of neutrophil degranulation and cytokine levels, including IL-1β, IL-6, IL-8, IL-10, TNF-α, and IFN-γ. BAL fluid is obtained with cannulation of the right middle lobe bronchus and infusion of 60 mL of saline utilizing three aliquots of 20 mL each. Each aliquot sample is obtained by brisk injection and slow withdrawal of fluid. MPO will be measured via a standard published method. Protein will be measured via a standard Bio-Rad protein assay kit II. A protocol for measuring the level of leukocyte surface expression by staining freshly isolated whole blood with a FITC-conjugated anti-porcine CD11b antibody has been developed. To assess leukocyte activation, leukocytes will be processed from heparinized blood which is either incubated immediately with a FITC-conjugated anti-porcine CD11b antibody or after a 30 minute exposure to PMA. PMA acts as a positive control/maximal stimulation condition. The antibody/blood mixture is then be fixed with a Becton-Dickenson's “Facs lyse” solution and the leukocyte population will then be analyzed for activated neutrophils, monocytes and eosinophils. The Becton-Dickenson's “Facs lyse” solution serves a three-fold purpose: it fixes the antibody-stained leukocytes for delayed flow cytometry analysis, it destroys the red blood cells, eliminating them from the analysis, and it separates the two granulocyte populations (neutrophils and eosinophils) from each other on the flow cytometer's forward and side scatter plots. Cells are then collected by centrifugation and resuspended for flow cytometry analysis for assessment of leukocyte activation of various cell populations. Early acute kidney injury (AKI) will be assessed by analyzing levels of neutrophil gelatinase-associated lipocalin (NGAL) in both the urine and the ultrafiltrate samples. Recent clinical data have clearly demonstrated that NGAL is an early biomarker for acute kidney injury. The amount of NGAL in the urine and serum is a highly specific and sensitive predictive marker of acute kidney injury. Relative NGAL levels will be determined by western blot using an antibody that cross reacts with porcine NGAL as indicated by the expected 25 kb band on western blot. Differences in NGAL levels should reflect the degree of AM in this animal model under sham and cell treated conditions. Serum elastase will be assessed as an indicator of peripheral blood neutrophil degranulation. Neutrophil elastase is released into the serum as neutrophils undergo degranulation. Under normal circumstances the serum elastase is inhibited by α1-antitrypsin, forming an elastase-α1-antitrypsin dimer. However, if the serum concentration of elastase exceeds the available inhibitor or α1-antitrypsin secretion by the liver is decreased due to hepatic injury, it may be possible to measure the activity of free elastase from the serum. Briefly, freshly isolated serum will be incubated in 100 μl reactions diluted in 1× Hank's buffered saline containing 50 μmol/L MeOSuc-AAPV-AMC (Biomol® International, Inc., Plymouth Meeting, Pa.) a fluorogenic substrate, incubated at 37° C. for 30 minutes. The conversion of fluorogenic substrate to detectable fluorescence will be measured on a microplate reader using the following settings: excitation wavelength, 380 nm; emission wavelength, 460 nm Controls will include samples with no added serum (negative control) and the addition of purified human neutrophil elastase (Biomol® International, Inc.).
Cardiovascular and biochemical data will be analyzed by repeated-measures analysis of variance (ANOVA). Plasma levels of various moieties, lung tissue and BAL MPO and elastase levels, leukocyte activation, NGAL fold over baseline, serum elastase and survival times will be compared utilizing Student's T-test, paired or non-paired as appropriate.
Contemplated Results
As suggested by the preliminary data detailed below, BRECS-d treated animals will have better cardiovascular indices, less pulmonary dysfunction and lung inflammation, and significantly extended survival times as compared to the sham group. BRECS-d treatment will act to inhibit leukocyte activation when compared to sham group.
As previously mentioned, a fully-freezable device has been designed and tested. This device is illustrated in
Prior to clinical application of the entire device, each component of the device is sterilized using, for example, ethylene oxide, gamma radiation, and/or autoclave. Thereafter, the device is frozen for four hours at negative 20 degrees Celsius, followed by an additional four hours at negative 80 degrees Celsius, followed by a transfer to liquid nitrogen vapor or liquid phase, as desired. For clinical application, the device can be thawed in a water bath, the DMSO may be flushed out, the device can be filled with media, and then the device can be connected to an extracorporeal circuit.
Flow and perfusion studies showed that the freezable BRECS-d of
Next, the freezable BRECS-d device of
This example describes the development of extracorporeal circuits (such as, for example, the ones shown in
In a pilot study, two different types of catheter were implanted to a sheep peritoneal dialysis model. Compared with an “Ash” Catheter, a double pigtail catheter is less likely to be clotted by omentum and fibrin. Tunneling the catheter as much as possible and/or wrapping external PD catheters with thicker tubing are helpful to better protect catheter from breaking or kinking.
Severe omental adhesion in autopsy was found to compart peritoneal cavity, probably due to reaction to omentectomy during catheter implantation. It was found that one-direction outflow obstruction is very common as a result of fibrin formation within the catheter. Therefore, high dose heparin at 2000 units in 10 mL saline is routinely added to each PD catheter once per day until commencement of PD. To reduce local inflammatory and traumatic response, omentectomy should not be performed when implanting PD catheter.
Optimization of Peritoneal Fluid Recycling with BRECS Unit
Catheters should be used at least 7 days post-op for better healing. In a pilot study, at least 2 liters fluid is needed to maintain appropriate inlet and outlet pressure, which is desirable to ensure PD recycling flow rate at about 100 mL/min A continuous supplement of fresh PD solution to the circuit is desirable in order to compensate for peritoneal absorption. This infusion rate should be at least 100 mL/hour. The first sheep was tested using a BRECS-d unit seeded with lamb kidney cells. Two peritoneal catheters were implanted into the peritoneal cavity, and antibiotic was routinely used to prevent infection. After 1 week post-op recovery, two liters peritoneal dialysate fluid with 1.5% dextrose (Deflex®, available from Fresenius Medical Care North America, Waltham, Mass.) were installed into the peritoneal cavity, with a continuous infusion rate of 100 to 200 mL/hour to supplement intra-peritoneal fluid absorption. The recycling PD circuit consists of two high flux polysulphone filters, whose membrane is both water- and solute-permeable, allowing for differentiated vectorial transport and metabolic and endocrine activity. Any bacterial contaminants with molecular weight larger than cut-off point of polysulphone membrane are held back and/or absorbed by the membrane, thus preventing development of peritonitis. A branch circuit at a desirable flow rate to sustain the nutritional and oxygen needs of the cells was initiated to maintain BRECS viability Immunoprotection of cultured cells in the BRECS unit is achieved concurrent with long-term functional performance as long as the filter membrane is intact. The BRECS is kept in a 37° C. temperature-controlled environment to ensure optimal functionality of the cells. The first sheep was maintained on recycling peritoneal dialysis for 10 days, validating feasibility of the recycling design.
Various compositions of PD solution have been used with this animal model. For example, peritoneal dialysate fluid with 1.5% dextrose solution (Deflex®; code 054-20201, Na 132 mEq/L, Ca 3.5 mEq/L, Mg 0.5 mEq/L, Cl 95 mEq/L, Lactate 40 mEq/L, Dextrose 1.5%, available from Fresenius Medical Care North America, Waltham, Mass.), 4.25% dextrose solution (Deflex®; code 054-20204, Na 132 mEq/L, Ca 3.5 mEq/L, Mg 0.5 mEq/L, Cl 95 mEq/L, Lactate 40 mEq/L, Dextrose 4.25%, available from Fresenius Medical Care North America, Waltham, Mass.), or Extraneal™ solution (Na 132 mEq/L, Ca 3.5 mEq/L, Mg 0.5 mEq/L, C; 96 mEq/L, Lactate 40 mEq/L, Dextrose 0%, Icodextrin 7.5%, available from Baxter International, Inc., Deerfield, Ill.) are used. For human subjects, peritoneal dialysate fluid with 2.5% dextrose solution (Deflex®; code 054-20202, Na 132 mEq/L, Ca 3.5 mEq/L, Mg 0.5 mEq/L, Cl 95 mEq/L, Lactate 40 mEq/L, Dextrose 2.5%, available from Fresenius Medical Care North America, Waltham, Mass.) can be used.
Although no potassium PD solution is used in this study, there are no significant changes in serum sodium and potassium concentration, probably compensated by the food intake of the sheep. There is a slight increase of calcium from 2.04 mEq/L at baseline to 2.31 mEq/L at day 7, but is still in the physiological range. Arterial blood gas was stable during the whole study period, reflecting a good ventilation and acid-base balance. The sheep's glucose metabolism is well-maintained even continuously exposed to high glucose environment, reflected by stable blood glucose level (69 mg/dl at day 1, 70 mg/dl at day 2). This data is displayed in
Prophylactic antibiotic was given intravenously. Data showed that combining nafcilin with gentamicin at an appropriate dose to balance the benefit of peritonitis prevention and the risk of pseudocolitis is beneficial. The first sheep developed peritonitis at day 4, presenting cloudy and bloody effluent. PD solution culture showed pseudomonas growth. The second sheep also caught a sepsis because of a pre-existing pneumonia and peritonitis. Thus, prevention and treatment of peritonitis in a prolonged treatment is desirable. Therefore, modifications (
Further results of the ex vivo testing of the circuits shown in
This research has set up systemic approaches of peritoneal dialysis catheter implantation, post-operation management, and a scheme of peritoneal fluid recycling circuit. The validity and feasibility of current designs will be further tested in a chronic uremic animal model. The peritoneal dialysis catheter implantation and post-operation management minimize the possibility of peritonitis and fibrin formation. Also, the peritoneal fluid recycling circuit is applicable to chronic uremic sheep models as well as other mammals, including humans, with BRECS unit viability maintained over time.
Several sham operations were performed with a sorbent delivery system from Renal Solutions, Inc. A sorbent regeneration system may be put into the circuit in parallel to the BRECS circuit. The used peritoneal dialysate will be regenerated at an interval of 4 times per day, with each regeneration lasting for about 1 hour, which is very close to inflow and drainage duration for one PD exchange in clinical practice. To ensure solute and water clearance, the speed of the infusion pump in the sorbent system will be programmed to ensure the final concentration of vital solute exactly the same as fresh dialysate.
Freezable BRECS-d Device with a Peritoneal Fluid Circuit
The freezable BRECS-d device, of
This example describes the further design and testing of the device illustrated in part in
A BRECS-h device will be fabricated that allows for cryopreservation/thawing under commercial manufacturing/clinical use—based conditions. The effects of cryopreservation on the BRECS-h viability and functionality will be determined under in vitro conditions. Additionally, conditions will be developed to optimize cell number in the BRECS-h unit, and BRECS-h maintenance parameters will be assessed to gain insight as to optimal conditions and for sustaining cell viability and functionality under simulated therapeutic circuit conditions. This research will further optimize the BRECS-h device in a pre-clinical large animal model of SIRS and sepsis, which can then be used to validate the device, for example, for use in humans.
The BRECS-h design allows for a simplified therapy circuit that uses a peripheral catheter access and a single blood pump system, such as the circuit shown in
The BRECS-h design utilizes cells grown on microcarriers or particles which are then resuspended in either a biomatrix 3-D gel or physiologic buffer, followed by seeding into the lumens of a hollow fiber bioreactor. The following method of deployment to the research/clinical site is contemplated. First, a tube of cryopreserved cells on microcarriers will either be mixed with a tube of pre-measured biomatrix or physiologic buffer. The cryopreserved cells and infusate medium would be stored at the research/clinical site or a central storage facility until needed. When needed, cryopreserved cells would be thawed, mixed with the biomatrix material or buffer, and syringe-loaded into the treatment device. Alternatively, the cryopreserved cells on microcarriers, already mixed with the biomatrix material or buffer (with both frozen), can be thawed and syringe-loaded into the treatment device. The cryopreserved tube could contain varying numbers of cells for dose-dependent therapy.
Hollow fibers have been successfully seeded with cells at near tissue density.
The initial BRECS-h studies were set up to determine optimal methods to obtain a hollow fiber loaded cartridge with 108 cells per BRECS-h unit. Because PPRECs are polarized epithelial cells that typically align to allow for apical and basolateral homogenous configuration, it was decided to first try seeding the cells in a 3-D matrix suspension, rather than cells alone, as with the HEP-G2 cells. Upon histologic assessment of cells resuspended in a 3-D matrix followed by seeding into a hollow fiber cartridge, it appeared desirable that the PPRECs use an attachment surface for optimal cell viability and expansion (data not shown). Study design was then shifted from free cells resuspended in a 3-D matrix, to cells grown on microcarriers, which are then either added to a 3-D matrix or physiologic buffer and seeded into hollow fibers.
Studies were undertaken to determine the feasibility of using a hollow fiber/microcarrier system to achieve a BRECS-h unit containing 1×108 cells. Cultispher® S microcarriers were used (available from HyClone, Inc. South Logan, Utah). They are a high thermally-stable, heavily cross-linked, porous gelatin bead.
PPRECs were seeded on gelatin-based Cultispher® microcarrier beads (available from HyClone, Inc.). These PPREC-seeded microcarriers were then grown in spinner flask culture for 21 days with samples taken for lactate production, oxygen consumption, and cell counts every day for two weeks and every other day for the last week.
At 5 days in culture, 150 μL of concentrated microcarriers were seeded into mini-hollow fiber devices containing 3 single hollow fibers per device (hollow fiber dimensions, 500 μm inner diameter, 660 μm outer diameter). Units were seeded under two conditions: suspended in a three dimensional alginate matrix and in renal cell culture media. The microcarrier seeded hollow fibers were maintained in culture for 5 days and then fixed for histologic examination.
To determine the feasibility of cryopreservation of the Cultispher® microcarriers, the following study was designed. Samples of microcarriers were taken from a spinner culture flask. Samples were taken for cell counts and DAPI staining. Microcarriers were frozen in one of three freezing solutions: CryoStor™ CS5, a cryopreservation medium containing 5% DSMO, (available from BioLife Solutions, Inc., Bothell, Wash.), 5% DMSO in renal cell media, or 10% DMSO in renal cell media. A 100 μL sample of concentrated microcarriers were placed in each of the respective freezing solutions with a final volume of 1 mL. The microcarriers were gently frozen by gradual temperature decrease in a Mr. Frosty freezing container (available from Nalge Nunc International, Rochester, N.Y.) filled with isopropyl alcohol for four hours at −20° C. and then transferred to liquid nitrogen for five days. The microcarriers were then quick thawed in a 37° C. water bath. Media was added drop-wise to a final volume of 10 mL. The microcarriers were incubated in 15 mL conical vials with 10 mL of media for 18 hours on a test tube rocker at 37° C. The microcarriers were then spun down, supernatant was removed, followed by trypsin digestion and trypan blue cell count to assess percent recovery vs. pre-thaw cell number estimates per vial and percent viability.
Table 2 summarizes the results from the three different cryopreservation solutions, with respect to percent recovery post-thaw of a reconstituted vial of cell/microcarriers vs. the estimated pre-thaw cell number per frozen vial of cell/microcarriers, the percent viable vs. dead cells (via trypan blue exclusion) in the post-thaw recovered cells and the percent viable cells in the post-thaw vial compared to the estimated cell number of the cell/microcarriers in the pre-thaw vial. The data show beneficial results from the cryo-protective agent, CryoStor™ CS5. The “% Post-Thaw Recovery of Estimated Pre-Thaw Cell Number” column shows the percent ratio of the number of cells that were frozen as assessed by trypsinizing a sample of the beads aliquoted in freezing solution and staining with trypan blue in comparison to the number counted when trypsinizing a sample of the beads under each condition post-thaw and staining with trypan blue. The number of both viable and non-viable cells are included in the cell count. The “% Death of Post-Thaw Recovery” column shows the percent ratio of the post-thaw dispersed cells that, when counted, are not viable based on trypan blue exclusion assay. The “% Viable Cells of Post-Thaw Recovered” column shows the percent ratio of post-thaw viable cells as a fraction of the total number of cells determined by cell dispersion from the microcarriers followed by cell counts (the data from the “% Post-Thaw Recovery of Estimated Pre-Thaw Cell Number” column), rather than the pre-thaw seeded cell number per gram of microcarrier. The “% Viable Cells of Estimated Pre-Freeze Cell Number” column shows the percent ratio of post-thaw viable cells as a function of the total number of cells prior to freezing, where total cells are the number of cells seeded per gram of microcarrier prior to the freeze/thaw process. The CS5 condition had 100% recovery of cells with almost 80% viability.
The assay discussed above in Example 2 regarding detection of leukocyte activation (under “Effect of PPREC-conditioned Media on Activated Leukocytes”) will be used to determine if conditioned media collected from the BRECS-h units after various manipulations and incubated with normal isolated porcine leukocytes under non-LPS and LPS stimulated conditions inhibit activation of leukocytes as compared to the non-manipulated BRECS-h media. Anti-porcine CD11b antibody staining of the leukocytes will be assessed and used as a biomarker index of the BRECS unit's ability to inhibit the activation of leukocytes. It is contemplated that the leukocytes, when exposed to PPREC-conditioned media, will be less activated at each condition than leukocytes not exposed to PPREC-conditioned media, as determined by CD11b flow analyzed studies.
Renal epithelial cells secrete low, but measurable levels of IL-8 under baseline conditions with significant increases post-LPS stimulation.
As shown above, the BRECS-h design allows for the device to be both miniaturized and the cells to be cryopreserved, followed by a quick thaw process and syringe loading, thus eliminating the shortcomings of the RAD design.
Development of a BRECS-h Prototype that can be Cryopreserved
The initial data gathered above will be further expanded.
Cell Growth and Seeding
PPRECs will be seeded and expanded on microcarrier beads. Optimal cryopreservation/thaw conditions will be established for the cell-seeded microcarriers. Thawed microcarriers will then either be resuspended in an alginate solution or in a physiologic buffer, seeded into a BRECS-h unit and, if alginate-based, gelled under increased calcium conditions. The BRECS-h system will also be tested in vitro to assess cryopreserved shelf life of the cell-seeded microcarriers 1 week, 1 month, 3 months and 6 months cryostorage.
The following parameters will be measured.
One consideration is the choice of cryoprotective agent. Cryoprotective agents such as glycerol and DMSO, which increase osmolarity, effectively dehydrate cells during the freezing process and thereby eliminate damaging ice crystal formation, will be considered for optimization. However, due to physical constraints, cells in solid organs, or in this case adhered to solid microcarrier structures, do not adjust as readily as cells in suspension to the abrupt changes in osmolarity inherent to standard cryopreservation protocols. Because microcarrier cultures have a high cell density, they should be transitioned quickly to avoid nutrient depletion; however, the diffusion of cryoprotectants is variable across solid structures and should be controlled to avoid mechanical stress to the surface cells while being allowed to permeate the microcarrier interior. A balance should be reached between these two factors.
Methods will be developed that optimize loading of cryopreserved PPRECs grown on microcarriers into hollow fibers post-thaw. It is contemplated that the cell-seeded microcarriers will be able to withstand cryopreservation with no detrimental effect on PPREC viability and function.
Optimization of Cell Number per BRECS-h Unit
Several methods will be tested to optimize both the number of cells per microcarrier and the loading capacity of microcarriers into the BRECS-h unit. Microcarrier cultures seeded and maintained under various conditions will be evaluated over time to obtain growth curves in culture to assess the number of cells per gram of microcarrier. The determined optimal growth parameters will then be used to test several loading protocols into the BRECS-h unit to develop a method that allows for maximal microcarriers per BRECS-h.
The following parameters will be measured.
It is expected that these methods will allow for BRECS-h units containing 108 to 109 cells. Further optimization of methods that allow for a slow, steady exchange of media with a flow through system that has a “catch trap,” so microcarriers are not removed with the spent media, may be desirable. Supplemental oxygen may also need to be supplied to the spinner flask system. Due to the potential of the hollow fiber packing density to make a non-homogeneous delivery of oxygen and nutrients to the encapsulated cell-seeded microcarriers, it is desirable to design hollow fibers with smaller outer diameter, which would be potted into the BRECS-h casing.
Determination of BRECS-h Maintenance Conditions that Sustain Optimal Cell Viability and Functionality
Conditions for flow rates, pressure gradients (both across the ECS to lumen path and along the pre- to post-ECS path), luminal buffer or alginate monomer composition, fiber length, and fiber density in the ECS housing will be assessed to determine which parameters support optimal oxygen and nutrient gradients to the cells grown on microcarriers and contained in the hollow fibers of the BRECS-h unit. Conditions will be simulated that will allow for the BRECS-h device to be assessed with respect to viability and functionality under parameters seen in an ex vivo extracorporeal circuit, such as the circuit of
Various conditions for flow rates, pressure gradients (both across the ECS to lumen path and along the pre- to post-ECS path), luminal buffer or alginate monomer composition, fiber composition, fiber outer diameter and wall thickness, fiber length, and fiber density in the ECS housing will be assessed to determine which parameters optimally support oxygen and nutrient gradients to the cells grown on the microcarriers contained in the hollow fibers of the BRECS-h unit. Conditions will be simulated that will allow for the BRECS-h to be assessed with respect to functionality under parameters seen in an ex vivo extracorporeal circuit with respect to oxygen and nutrient delivery and pressure gradients. In order to more closely match the clinical circuit setting, heparinized bovine blood will be used in the in vitro test circuit for some of the studies in this component of the research plan. Functional assessment criteria will include lactate production, oxygen depletion, glucose consumption rates, histologic assessment, leukocyte conditioned media assay and IL-8 baseline and LPS stimulated secretion rates. Many of these studies (luminal buffer or alginate monomer composition, fiber composition, fiber outer diameter and wall thickness) will initially be performed in mini-hollow fiber units to test optimal parameters. Once these have been determined, then the BRECS-h units will be used for testing final parameters for flow rates, pressure gradients and fiber packing density. 8 BRECS-h units will be tested for each of these parameters with 8 BRECS-h unit being tested with the final optimal combination ((3 test parameters×8 BRECS-h per parameter)+8 BRECS-h for optimal combination=32 BRECS-h units).
It is contemplated that the cells in the BRECS-h will prove both viable and functional with respect to the assessment test parameters. Conditions for flow through the BRECS-h ECS will allow for adequate nutrient and oxygen delivery over the anticipated therapeutic application time period. Hollow fiber length and/or hollow fiber packing density inside the ECS may be varied. A membrane oxygenator may be used in the pre-BRECS-h section of a circuit to enhance further delivering oxygen to the cells. Such an optimized BRECS-h unit may be tested in a pre-clinical large animal model of sepsis and SIRS that would target the septic shock and SIRS patient population. This model would confirm the utility of the BRECS-h and validate its use in humans.
This example describes a protocol for coating substrates according to the invention, such as the Nb-coated carbon disks of
Collagen Coating of Carbon Nb Disks
The technique for passive coupling of proteins to the Dynamatrix carbon Nb disk matrices recommended by Dynamatrix, the manufacturer, is a one-sided static coating technique. The protocol from the manufacturer calls for disks to be coated by placing full tuning fork assemblies, with four disk matrices, each into tissue culture plates and loading 150 μl of protein solution onto each individual disk. A concentration of 10 μg/mL is recommended by the manufacturer for the protein solution, but the protocol states that other concentrations may be appropriate depending on the specific protein and application. After the solution is added, there is an incubation step of one hour at room temperature to allow for protein adsorption onto the matrices. Excess fluid is then aspirated, and the matrix units are dried at room temperature in a sterile hood for one hour prior to cell seeding.
In order to optimize collagen IV coating of disk matrices, the distribution of collagen IV on the disks achieved by static coating was compared to an alternative coating technique in which collagen solution was loaded onto each side of the disk, and centrifugal force was applied after loading. The applied force was intended to push collagen solution into the center of the matrix in an attempt to produce more uniform collagen coverage throughout the matrix than the coverage achieved by the manufacturer's static loading protocol. The distribution of collagen through the disk matrices was evaluated qualitatively by assessing the signal of fluorescent-conjugated collagen IV on the top, bottom, and cross-sectional middle of disks coated by both single-sided static and double-sided centrifugation techniques.
The manufacturer's static seeding technique resulted in heavy coverage of the top of the carbon Nb disk, moderate coverage of the bottom of the disk, and minimal coverage of the middle of the disk. The bottom of the disk showed a strong collagen signal around the edge which rapidly decreased toward the center of the disk, implying that most collagen solution passed around the disk instead of through it to fill the inner matrix. An amended static technique in which collagen was loaded to both sides of the disk resulted in heavy coverage of the top and bottom, but still resulted in minimal coating in the middle of the disk. Since collagen coating occurs by protein adsorption, the lack of coating in the middle of the disks is probably due to passive loading of the collagen to the top of the disk, allowing protein solution to simply pass around the edges of the disk. Without any force applied to push collagen through the disk, the center of the disk does not have adequate contact with collagen to allow for adsorption.
Double-sided coating with force applied by a one-minute 14.84×g (300 rpm) centrifuge spin cycle on each side resulted in similar collagen coating of the top of the carbon Nb disk and improved coating of the middle and bottom of the disk compared to static coating. The applied centrifugal force improves the distribution of collagen solution throughout the disk and results in more uniform collagen adsorption by ensuring that collagen passes into the center of the disks where it can be adsorbed during incubation. The coating of the bottom surface of the disk is uniform as a result of loading collagen to both sides, in contrast to the non-uniform coating of the bottom surface of the disk loaded on only one side.
From the results of these collagen coating studies, the protocol for coating carbon Nb disks with collagen IV has been optimized to a double-sided centrifugation coating technique. Collagen solution is prepared at a concentration between about 0.75 ng and about 1 μg collagen IV per cm2 of disk matrix to be coated. Prior to coating, all disks are removed from tuning fork assemblies and placed into separate wells of a 48-well cell culture plate. After collagen solution is adjusted to a pH between 7.0 and 7.5, 75 μl of collagen solution is loaded to the top of each disk. Immediately following loading, the culture plate is spun at 300 rpm for one minute. The disks are then incubated in a laminar flow hood at room temperature for one hour to allow for collagen adsorption. Each disk is then turned over, and 75 μl of collagen solution is added to the second side. The culture plate is again spun at 14×g (300 rpm) for one minute, followed by a one-hour incubation at room temperature to allow for adsorption. The disks are then transferred to a new culture plate for cell seeding.
Cell Seeding onto Carbon Nb Disks
The technique for cell seeding onto the Dynamatrix carbon Nb disk matrices recommended by the manufacturer is a one-sided static seeding technique similar to the collagen coating protocol described above. The protocol calls for disks to be seeded by placing full tuning fork assemblies with four disk matrices each into tissue culture plates. A 75 μl cell suspension is then added drop-wise to each individual disk. The total number of cells to be seeded onto each unit is recommended within the range 2.5×105 to 5.0×105 cells, but the protocol states that other seeded cell totals may be appropriate depending on user need and experience. After the cell suspension is loaded, an incubation step of 2-4 hours in a humidified incubator allows for adhesion of cells onto the matrices.
To optimize the seeding of renal epithelial cells onto the carbon Nb disks, distribution of cells throughout the matrices achieved by manufacturer-recommended protocol was compared to an amended protocol in which cells were statically seeded to both sides of the disk. A disk seeded following manufacturer protocol was seeded on one side and incubated for 3 hours, while a disk seeded by the amended protocol was seeded on both sides with a 1.5 hour incubation following seeding onto each side. Distribution of cells throughout the disk matrices was evaluated qualitatively by assessing the signal of fluorescent cell stain on the top surface, bottom surface, and cross-sectional middle of disks comparing manufacturer's standard protocol and the amended double-sided protocol.
The amended double-sided seeding technique resulted in an improved, more uniform distribution of cells throughout the carbon Nb disk as compared to the manufacturer's one-sided seeding technique. Cell coverage was especially improved in the middle and on the bottom surface of the two-sided disk.
One-sided seeding demonstrated heavy cell coverage of the top surface of the disk, moderate coverage of the middle of the disk, and minimal coverage of the bottom surface of the disk. The difference in cell coverage of the top and bottom surfaces results in disks that are sided and could affect the performance of the disks in the BRECS-d units. The one-sided seeding disk shows that the renal cells are able to pass through the disk pores and attach to the interior without external applied force, but loading cells from only one side of the disk showed incomplete cell coverage.
Double-sided seeding produced heavy coverage of the top surface, bottom surface, and middle of the disk. Compared to one-sided seeding, cell coverage was much more uniform throughout the double-sided disk. Loading cells on both sides eliminates the previous issues as cell coverage is more uniform between bottom and top surfaces. Double-sided static seeding still allows cells to pass into and attach to the interior of the disk, and the resulting cell uniformity makes this technique a more effective option than single-sided loading.
From the results of the cell seeding study, the optimized protocol for seeding renal epithelial cells onto carbon Nb disks coated with collagen IV is a double-sided static seeding technique. Prior to seeding, carbon Nb disks are coated with collagen IV by the optimized coating protocol discussed above. The coated disks are placed into separate wells of a 48-well cell culture plate for seeding. A cell suspension of renal epithelial cells in a customized UltraMDCK media (available from Lonza Group Ltd., Basel, Switzerland, catalog number 2042042) is prepared with a concentration of 1×106 cells per 75 μl of media. 1 L of UltraMDCK media is mixed with 1 ml of a 500× Insulin, Transferrin, Ethanolamine, and Selenium (ITES) mixture for a final concentration of 0.5× (available from Lonza Group Ltd., Basel, Switzerland, catalog number 17-8392), 3 μl Triiodothyronine (T3) of 2×10−6M for a final concentration of 6×10−12M (available from Sigma-Aldrich, Inc., St. Louis, Mo., catalog number T-6397), 500 μl Epidermal Growth Factor (hEGF) of 2×10−5M for a final concentration of 1×10−8M (available from R&D Systems®, Minneapolis, Minn., catalog number 236-EG), and 100 μl Retinoic acid (RA) of 1×10−3M for a final concentration of 1×10−7M (available from Sigma-Aldrich, Inc., St. Louis, Mo., catalog number R-2625). Alternatively, the concentrations of the additives to the UltraMDCK mixture can be of various ranges. For example, the final concentration of the ITES mixture can be of a range from about 0.0 to about 0.5×, the final concentration of T3 can be of a range from about 0.0M to about 6×10−12M, the final concentration of hEGF can be of a range from about 0.0M to about 1×10−8M, and the final concentration of RA can be of a range from about 0.0M to about 1×10−7M. UltraMDCK media is used for renal cells. Alternatively, other media are contemplated for other types of cells, such as other eukaryotic cells, such as other mammalian cells. A 75 μl cell suspension is then added to each disk, followed by 1.5 hours of incubation at 37° C. and 5% CO2 to allow for adhesion of cells to the matrices. Each disk is then turned over, and an additional 75 μl of cell suspension is added to the second side. A second 1.5 hour incubation period at 37° C. and 5% CO2 is carried out to allow for cell attachment.
Post-Seeding Culture of Carbon Nb Disks
Post-seeding culture of Dynamatrix carbon Nb disks by manufacturer protocol is performed in the Starwheel system. The system consists of a media reservoir which has a Starwheel spinner connected to the cap. The spinner has eight arms into which tuning fork assemblies having cell-seeded disks are placed. The spinner is magnetic and spins at a set rpm when the Starwheel system is placed on a stir plate. Manufacturer protocol calls for seeded tuning fork assemblies to be placed into spinner arms in a sterile environment. The cap is then secured onto the media reservoir. The reservoir is filled with 175 mL of media through a side arm, and the system is placed on CellGro spinner plates. Finally, the two side arm caps are loosened to allow for gas exchange during culture, and the spinner plate is set to the desired rpm. Culture rpm recommended by the manufacturer is 2 rpm, and the protocol warns against exceeding 10 rpm.
To optimize post-seeding culture of carbon Nb disks seeded with renal epithelial cells and to provide for a more robust manufacturing process, culture in the Starwheel system was compared to culture in a BRECS-d unit (such as the one shown in
Perfusion in the BRECS-d systems resulted in better cell growth and more effective maintenance of viable cells on the carbon Nb disks than the Starwheel culture systems. After two weeks in culture, both the 25 mL/min and 50 mL/min BRECS disks showed improved cell coverage compared to either of the Starwheel system conditions. There was no noticeable difference in cell coverage between the two BRECS-d perfusion systems, but both had significantly higher fluorescent cell signal than both of the Starwheel systems.
From the results of the post-seeding culture study, an optimized protocol has been developed which does not need to use the Starwheel system. Instead of culturing seeded disks in the Starwheel system prior to use in the BRECS-d system, seeded disk matrices in tuning fork assemblies are placed directly into the BRECS unit. Culture in the BRECS-d is performed by the perfusion of UltraMDCK media through the unit at rate of 25 to 50 mL/min Alternatively, a perfusion rate of between about 10 mL/min and about 200 mL/min can be used.
This example describes the development of a cell therapy system implemented in two large animal applications, an anephric sheep model of uremia and a porcine model of sepsis. The goal was to evaluate the functionality of cell-seeded BRECS-d units in pre-clinical acute and chronic therapy applications. More particularly, the chronic uremic sheep model was used to test the BRECS-d device of
As described in Example 4, a twenty-four hour continuous flow peritoneal dialysis (CFPD) protocol was developed using normal sheep. This protocol utilized the LREC-seeded BRECS-d device of
Results
CFPD was started and maintained successfully without PD catheter-associated complications or peritonitis in 10 animals (n=3 for normal sheep and n=7 for uremic sheep; average values of the data collected for the CFPD studies were calculated based on the number of animals that remained in the study at a particular time point). Further, CFPD provided a flow rate of ≧80 mL/min in the main circuit and a flow rate of 50 mL/min within each BRECS-d unit without interruption. This was achieved while maintaining a intra-peritoneal dwelling volume of 2 L, a dialysate infusion rate of 400 to 450 mL/hr, and a dialysate drainage rate of 450 to 470 mL/hr. Given this difference between the dialysate infusion and drainage rate, the net change in intraperitoneal fluid volume, otherwise known as the net ultrafiltration volume, in each uremic animal was estimated to be between about 480 mL/day to about 1200 mL/day.
During the study period, viability of cells in BRECS-d units was assessed by measuring O2 consumption. As shown in
Neither interruption of circuit flow nor cell death occurred during the study period. The volume and hemodynamic state and chemistry profiles were adequately controlled in both models. Uremic control by CFPD measured by small solute uremic parameters was within acceptable ranges in the uremic sheep.
The effect of BRECS-d cellular therapy on the inflammatory state of uremic sheep was assessed (using the device of
The utility of cell-seeded BRECS-d units in an acute therapy application was studied. In this regard, PPREC-seeded BRECS-d units (as shown in
Cardiovascular parameters were measured and samples for biochemical analysis were obtained at various time intervals during the experiment. PPREC-seeded BRECS-d treated animals showed overall improvement in cardiovascular parameters, with a 27% improvement in cardiac output. Further, BRECS-d therapy increased average survival time from 6.7 hours in the non-treated sham group to 11.3 hours in the BRECS-d treated group. (Of note, 3 BRECS-d treated animals were terminated at the predetermined 12 hour end point rather than expiring due to sepsis).
Experiments are underway to confirm beneficial time frames for application of frozen and thawed BRECS-d devices. Preliminary studies will initially place a thawed unit immediately on the pig. The time to application also will be varied. Evaluation of these devices will confirm beneficial lag times before placement and use in clinical settings.
The peritoneal dialysis treatment circuit of
This example describes the durability of PPREC- and HREC-seeded BRECS-d units of
Long term in vitro culture studies were conducted to determine the durability of PPREC- and HREC-seeded cell devices. Cell longevity and functionality in BRECS-d devices are relevant to their targeted chronic therapeutic applications. As such, the viability and metabolic activity of the cells in BRECS-d units were assessed.
To assess whether cells cultured in BRECS-d devices maintained differentiated renal characteristics over time, HREC-seeded BRECS-d units (as shown in
Results from these studies indicate that HRECs cultured in BRECS-d devices maintained differentiated renal characteristics for up to 90 days. As shown in
As an additional test of the functionality and differentiated state of HRECs maintained in BRECS-d units of
As a valuable component to clinical application of a cell therapy device, cryopreservation of BRECS-d units and storage options were evaluated. Liquid nitrogen vapor phase cryopreservation was assessed to determine material integrity under freezing and thawing conditions. Protocols were evaluated to investigate thermodynamic control in BRECS-d units. In these studies, BRECS-d units, as shown in
Thawing time is relevant to cell viability and clinical implementation. The initial thaw time for an entire BRECS-d unit was over 45 minutes, and sometimes lasted up to 1.5 hours. To further establish a thawing protocol, twenty six trials were conducted to develop the protocol described below. Each trial involved a single acellular device of
PPREC-seeded BRECS-d units (shown in
PPREC-seeded BRECS-d units were thawed according to the above-described protocol for acellular devices, resulting in a transition from −140° C. in the liquid nitrogen vapor phase to a 37° C. operating temperature within 20 minutes. The central core of each BRECS-d unit, which contained the cell-seeded disks, remained frozen during the initial 12 minute incubation at 90° C. The cells were subsequently thawed by perfusion with 37° C. media, which rapidly replaced the cryopreservation buffer.
Many clinical facilities have −80° C. freezers but are less likely to possess on-site liquid nitrogen storage. Therefore, additional testing was conducted to confirm that BRECS-d units (shown in
Lastly, studies were done to assess shipment to the clinical site. Dry ice was used to maintain temperature during shipping studies. A study using thermocouples placed inside a cryopreserved acellular BRECS-d unit of
These studies detail the design, fabrication and successful testing of a compact, cryopreservable BRECS-d unit. This innovative, tissue-engineered device provides enabling technology for use in acute and chronic disease applications
The entire content of each patent and non-patent document disclosed herein is expressly incorporated herein by reference for all purposes. Moreover, U.S. patent application publication No. 2007/0269489 is expressly incorporated herein by reference for all purposes.
The invention may be embodied in other specific forms without departing from the spirit or essential characteristics thereof. The present embodiments are therefore to be considered in all respects as illustrative and not restrictive.
This application is a continuation of the International Patent Application PCT/US09/53516, filed Aug. 12, 2009, which claims priority to and the benefit of U.S. Provisional Patent Application No. 61/189,119, filed Aug. 15, 2008, the entire disclosures of each of which are incorporated herein by reference.
This invention was made with government support under grant number W81XWH-05-2-0010, and modification numbers P00004 and P00005, awarded by the U.S. Army Medical Research and Materiel Command, and grant number 2R44DK074289-03, awarded by the National Institute of Diabetes and Digestive and Kidney Diseases. The government has certain rights in the invention.
| Number | Date | Country | |
|---|---|---|---|
| 61189119 | Aug 2008 | US |
| Number | Date | Country | |
|---|---|---|---|
| Parent | PCT/US2009/053516 | Aug 2009 | US |
| Child | 13027481 | US |
