EX-10 5 nih.txt PREPARED BY: MHUEBOTTER@HOTMAIL.COM EXHIBIT BIOGRAPHICAL SKETCHES AND BIBLIOGRAPHY DR. NINA LAMBA (Principal Investigator) graduated in 1991 with a B.Sc. (Hons.) degree in Applied Chemistry from Aston University, Birmingham, U.K. She was awarded a doctorate in Bioengineering in 1994 from the University of Strathclyde, U.K. Her doctoral thesis "Blood-biomaterial interactions: Application of a parallel plate flow cell to study blood responses in vitro" was conducted under the supervision of Professor J. M. Courtney. Following graduation, she joined the Department of Chemical Engineering at the University of Delaware as a post-doctoral research fellow and Research Associate, working with Dr. Stuart L. Cooper. Her research activities included studies of cellular adhesion to polyurethanes, the interrelationships between infection and inflammation associated with biomaterials and aspects of tissue engineering. She joined Compact Membrane Systems, Inc. last year as Director of the Biomedical Research Division. She has made significant contributions to the biomedical literature, and selected publications are listed below. She is also a member of the Society for Biomaterials, the American Society for Artificial Internal Organs, and of the Royal Society of Chemistry. N.M.K. Lamba, K.A. Woodhouse, S.L. Cooper. "Polyurethanes in Biomedical Applications," CRC Press, 1998, pp. 277. N.M.K. Lamba, J.D.S. Gaylor, J.M. Courtney, G.D.O. Lowe. "Complement activation by cellulose: Investigation of the effects of time, area, flow rate, shear rate and temperature on C3a generation in vitro, using a parallel plate flow cell."J. Mater. Sci ., Mater. Med. (1998) 9(7): 409-414. N.M.K. Lamba, J.A. Baumgartner, S.L. Cooper. "Cell-Synthetic Surface Interactions" In: Frontiers in Tissue Engineering. Eds. C.W. Patrick Jr., et al. Elsevier Science Publishers. 1998, p. 121-137. N.M.K. Lamba, S.L. Cooper. "Covalent grafting of RGD peptides to synthetic surfaces." In: Tissue Engineering of Prosthetic Vascular Grafts. Eds. P. Zilla and H.P. Greisler. R.G. Landes Co. 1999, p. 553-559. N.M.K. Lamba, J.D.S. Gaylor, J.M. Courtney, G.D.O. Lowe. "Effect of heparin on the blood response to blood-contacting biomaterials in vitro." Biomaterials. (2000) 21(1): 89-96. DR. STUART NEMSER - Dr. Nemser received his BS. in 1966, MS. in 1968, and Ph.D. in 1972 all from MIT in Chemical Engineering. Between 1972 and 1974, he worked for Millipore Corporation (Bedford, Mass.) and developed their line of Thin Film Composites (TFCs), ultrafiltration membranes via interfacial membranes formation ("Pellicon" line) as well as developing their supported microfiltration membranes for pleated cartridges. Dr. Nemser worked for DuPont from 1974 until his retirement in May 1993 to found Compact Membrane Systems, Inc. During his 19 years at DuPont, Dr. Nemser held numerous research, business and manufacturing positions, including Technical Superintendent of DuPont's "Permasep" reverse osmosis membrane business. As New Business Development Manager for E. I. DuPont, Dr. Nemser identified and patented the family of PDD homopolymers and copolymers for various separations. In his current position as President of CMS, Dr. Nemser is focused on commercializing the family of PDD polymers in the patented membrane areas. Dr. Nemser has been involved in a number of NIH funded projects, including IVOX and ECMO, membrane oxygenators for bioreactors, ozonation of drinking water and dental unit water lines, and portable O2 enriching systems for lung disorders. These programs have provided Dr. Nemser with excellent experience in enhancing oxygenation delivery to fluids through their novel bubbleless oxygen delivery nonporous membranes. Key publications include: S.M. Nemser, "Controlled Porosity Reverse Osmosis Membranes," MIT Doctoral Thesis, June 2, 1972 S.M. Nemser and I. C. Roman, "Perfluorodioxole Membranes," U.S. Patent 5,051,114, (Sept, 24, 1991) S.M. Nemser, "Air-intake Systems for Mobile Engines," U.S. Patent 5, 147,417, Sept. 45, 1992 S.M. Nemser, "Air-intake systems for Residential Furnaces:, U.S. Patent 5, 053, 059, Oct. 1, 1991 DR. HUGO O. JAUREGUI - Dr. Jauregui received his MD from the University of Buenos Aires School of Medicine and his Ph.D. in Pathology from Duke University. Dr. Jauregui has two decades experience in the development of cell-based liver assist devices. In 1979, his laboratory was the first to explore the culture of rat hepatocytes in perfused hollow fiber bioreactors. Numerous studies aimed to identify hepatocyte culture requirements and optimize culture in these devices ensued. These perfusion bioreactors were later modified (in collaboration with Dr. Barry Solomon of Amicon) to create liver assist device prototypes for testing in experimental models of hepatic encephalopathy. Dr. Jaregui's group characterized the galactosamine (gal) intoxicated rabbit model which nearly mimics hepatic encephalopathy in humans. Studies to scale-up rat-rabbit hepatocyte isolation procedures led to ex vivo porcine hepatocyte procurement by collagenase digestion. Further activities of Dr. Jauregui's group were directed toward storage and transport requirements of porcine hepatocyte preparations. Proprietary methodology was developed for the frozen storage of porcine hepatocytes which allows recovery of ~80% viable cells which maintain ~75% of their initial P450 metabolic function. Additional activities explored methods of culture which would allow scale-up of porcine hepatocyte populations from monolayer. Data generated from these porcine hepatocyte studies helped guide the development of the methodologies currently used in the Circe Biomedical ELAD system today. Currently, Dr. Jauregui is the President of MultiCell Associates, Inc., a company dedicated to the development and commercialization of cells and cell lines for medical use, with a focus on developing immortalized human hepatic cell lines which have retained their primary functionality. Some publications and patents include: Jauregui HO, C Mullon, P Press, D Trenkler, S Naik, H Santangini, T Muller, B Solomon. 1995. In Vivo evaluation of a hollow fiber liver assist device. Hepatology 21:460-469. Jauregui HO, N Roy Chowdury, J Roy Chowdhury. 1996. Use of mammalian liver cells for artificial liver support. Cell Trans 5,3:353-367. Jauregui HO, S Naik, H Santangini, D Trenkler, C J-P Mullon. 1997. The use of microcarrier-roller bottle culture for large scale production of porcine hepatocytes. Tissue Eng 3(1):17-25. Jauregui HO. 1997. The technology of biological extracorporeal liver assist devices: From infancy to adolescence. Artif Org. 4721(11). Naik S, H Santangini, D Trenkler, C J-P Mullon, B Solomon, HO Jauregui. 1997. Functional recovery of porcine hepatocytes after hypothermic or cryogenic preservation for liver support systems. Cell Trans, 6(5):447-454. Liu J, P Jing, S Naik, H Santangini, D Trenkler, N Thompson, A Rifai, JR Chowdhury, HO Jauregui. 1998. Characterization and evaluation of detoxification functions of a Nontumorigenic Immortalized Hepatocyte Cell Line (HepLiu). Submitted to Cell Trans. Patents: Patent #5,043,260. Perfusion Device with Hepatocytes, Patent #4,795,459. Implantable Prosthetic Device (Endothelial), Dual Fiber Bioreactor. (Patent allowed), Isolation and Culture of Porcine Hepatocytes ( USSN 08/5411,462 filing date 12/22/95), Cryopreserved Hepatocytes and Process for Doing Same. (Patent Pending), Immortalized Hepatocytes. (USSN 08/611, 171 (allowed). A. SPECIFIC AIMS The long range objective of the proposed R & D project is to enhance the function of the bioartificial liver (BAL) device by improving the external bubbleless oxygenation system used for hepatocyte culture. More specifically, we plan to fabricate and evaluate a new bubbleless external hollow fiber oxygenator for use in the Sybiol(R) Bio-Liver Device developed by Exten Industries. 2 A unique, non-porous, highly gas permeable amorphous perfluorocarbon polymer (CMS-7) will be applied as a very thin membrane on microporous polypropylene hollow fibers. CMS-7 will provide excellent oxygen transport with minimal deposition and fouling. The non-porous perfluoropolymer layer will provide minimal resistance to O2/CO2 flow while stabilizing the system's long term performance. Non-wetting porous hollow fiber membranes (HF) have good initial O2/CO2 transport, but deteriorate over time due to fluid infiltration into the pores and breakthrough of the perfusate into the gas side of the oxygenator (wet out). By coating porous HF with a thin continuous layer of our high flux and non-porous perfluoropolymer film, the initial flux will be comparable to other porous HF (in fact, coating of porous HF actually increases oxygen flux-see Tables 1, 2; Fig. 3), with minimal degradation over time. With increased, long-term access to oxygen, hepatocyte viability and detoxification function is expected to improve. Prolonged hepatocyte function will provide better treatment for patients with acute or chronic hepatic failure. Specific aims include: 1. Fabricate CMS oxygenators using polypropylene hollow fiber modules coated with the CMS-7 polymer. 2. Compare a number of modules with and without the CMS membrane by analyzing mass transfer of oxygen into water and hepatocyte media, varying conditions to optimize performance. 3. In vitro, analyze oxygen mass transfer with recirculating media containing hepatocytes. Compare CMS oxygenators to those without the CMS membrane (controls), and monitor cell viability relative to oxygen mass transfer. 4. Integrate the oxygenator and recirculating cell culture with the hollow fiber module used as a blood/cell culture interface to monitor cell viability and detoxification function in vitro, comparing controls to coated modules. 5. Plan Phase II, which will include comparison of primary and immortalized hepatocytes, and in vivo animal trials. Food and Drug Administration (FDA) approval will be secured for clinical trials for device safety. B. SIGNIFICANCE One in ten Americans has some form of liver disease. Currently, the treatment of choice for patients with acute hepatic failure (AHF) and chronic hepatic failure is liver transplant. AHF is brought on by Hepatitis or an overdose of drugs such as acetaminophen, and can lead to encephalopathy (HE) (mortality of 50-80%)(1). However, in the United States, there are only about 3000 donor livers available annually, while approximately 30,000 patients die from liver failure (2). The American Liver Foundation states that liver failure is the fourth leading cause of death in the United States (40,000 deaths in 1996). According to Algenix, Inc. of Minneapolis, a company currently developing a BAL, the market potential for the BAL industry exceeds $700 million in the United States, and may reach $1.4 billion worldwide (23). Transplantation procedures are expensive, and carry risks increased by AHF. If a patient suffering from AHF could receive temporary liver support, the native liver would have time to regenerate. Patients with advanced liver disease have been known to recover once regeneration begins, often with no residual disease after recovery (3). In fact, using BAL while awaiting organ transplant yields better survival rates, less frequent graft failure, and earlier hospital discharge than those patients whose health deteriorates prior to transplantation due to poor liver function (3). A bioartificial liver assist device (BAL) can serve as a "bridge" during reversible AHF, allowing the native liver to regenerate, and can prevent multisystem failure as a result of the accumulation of metabolic byproducts. For patients with chronic liver diseases, BAL could also be used for periodic liver "dialysis," or detoxification. The ideal BAL would mimic normal liver function as closely as possible for as long as possible while requiring minimal investment. Many of the BALs being developed today have focused on providing support for a short period of time (six hours) as a bridge until transplantation, but for the native liver to regenerate, the BAL must operate for a longer duration (at least ten days) (21). In addition, pharmaceutical companies can use BALs for drug metabolism evaluations that would be safe and inexpensive. Our oxygenation technology is non-fouling and resists wet out, indicating that it would be advantageous for long term BAL operation. Adequate oxygenation is essential in BAL devices to obtain maximum cell viability and function. CMS membrane technology will improve oxygenation of the hepatocyte cell culture, allowing better control of oxygen flux while resisting "wet out", to provide maximum hepatocyte functionality. 3 1. Variations of BAL ----------------- BALs vary in the choice of perfusate, the choice of cell, and the arrangement of hepatocytes within the bioreactor. Either the patient's blood or plasma is perfused through a bioreactor containing the hepatocytes. Clinically tested BALs have employed transformed cell lines (e.g. C3A, a cell line derived from human hepablastoma) or freshly isolated porcine hepatocytes. Since human hepatocytes are of limited availability, rabbit, rat, and pig hepatocytes have been successfully used in experimental BALs. Porcine hepatocytes can be procured in adequate numbers, screened for contamination by other cell types, and cryopreserved so a continuous supply is assured (4). Immortalized porcine hepatocytes have also been developed (26) and may be used in Phase II of this project. Different polymer materials and/or configurations of the man-made components used in BALs have undergone the most advances. Hepatocytes have been used in suspension or immobilized on a nonspecific material, and have been cultured among three dimensional nonwoven polyester fabric, with HF oxygenation (5-7, 9). Some of the most promising and prevalent design technology has involved hollow fiber membranes (HF). Hepatocytes are cultured in a variety of arrangements, including in suspension (40) within a gel bed inside the HF (8), outside the HF in suspension, in cell aggregates, or bound to microcarriers (7, 8), in adhesion outside of HF (3, 8), or among mats of woven capillary membranes (7, 8). For this Phase I project, CMS has chosen to work with Exten Industries to develop an in-series HF oxygenator for use with their Sybiol(R) synthetic bio-liver device, discussed below. 2. The Sybiol(R) Bio-Liver Device ------------------------------ The Sybiol(R) synthetic bio-liver device is now being developed by Exten Industries Corporation. The Sybiol(R) bioartificial-liver circulates patients' blood through a replaceable cartridge (hollow fiber module), where it is interfaced with constantly recirculating viable hepatocytes. Toxins and enzymes diffuse across the membrane, and are metabolized by the recirculating hepatocytes. The expectation behind all types of bioartificial liver devices is that by providing liver support with the appropriate metabolic, detoxification and biosynthetic functions, the native, ailing liver could be given an opportunity to recover from shock, disease, poisoning, or drug overdoses. In cases of acute hepatic failure, it is envisaged that bioartificial liver support could allow regeneration to the degree that no transplant would be required, reducing the demand for the already limited supply of donated organs. Furthermore, such a device could also help liver transplant patients prepare for surgery and provide post-transplant support. The Sybiol(R) technology circulates blood extracorporeally, as depicted in the diagram below (Figure 1). This device may have significant advantages in methodology and mechanical design that allow tangible benefits in treatment and cost over competitive technologies currently in development. Although it has been shown that hepatocyte adhesion improves liver function in vitro (5,6,7), attachment to microcarriers or membranes does not necessarily determine liver function. Reseachers working on the Sybiol(R) liver device demonstrated that hepatoctyes remained viable in suspension without prior immobilization onto microcarriers (42). A major advantage of recirculating hepatocytes rather than isolating them in the extracapillary space, is that there is a less stringent limitation on the number of cells that can be included in the circuit. A recirculating loop also offers the ability to replenish the supply of hepatocytes as needed, and allows closer monitoring and control over variables such as temperature, pH, etc. Hepatocytes in BALs that have been clinically tested lose function in 6-8 hours. By increasing the number of other hepatocytes and adding the ability to replenish the hepatocyte supply with new cells, the Sybiol(R) synthetic bio-liver device should prolong performance. A recirculating system also has potential to be used as liver "dialysis" treatment for either chronic liver disease stemming from alcoholism or hepatitis, and/or to treat ingestion or overdose of toxic substances. Other advantages include lower cost, fast change-over between patients, and a much longer treatment cycle. The design of this technology also minimizes the possibility of immune reactions. 4 An additional role for this recirculating bioartificial liver is in the evaluation of drug metabolism. Therapeutic drugs introduced into the body are metabolized by the liver, and may have a variety of beneficial as well as detrimental effects. It is, therefore, mandatory that developers of new drugs thoroughly evaluate this process. A design such as this will allow for the sampling of not only the media used to suspend the cells, but for the easy sampling of hepatocytes in order to evaluate any phenotypic changes. The availability of a laboratory device able to accurately simulate the behavior of the human liver will provide a significant alternative to animal testing for the pharmaceutical industry. A CMS oxygenator will reduce fouling and "wet out" compared to a HF oxygenator without the membrane, thus prolonging oxygenator performance. A long term oxygenator will be necessary if the Sybiol(R) is to deliver liver treatment for a sufficient amount of time. Eliminating the need to change the oxygenator also reduces the likelihood of introducing pathogens into the heptocyte suspension. Currently, oxygenation of the hepatocytes is achieved by running the suspension through a 28ft length of silicone rubber tubing, which is wound around a mandrel. O2 and CO2 diffuse passively through the tubing wall, and this is the source of O2 for the cells. Replacing this coil with a membrane oxygenator will allow increases in the rates of oxygen and carbon dioxide transfer and better control over oxygen transfer through the increase in membrane permeability and the use of enriched gas feedstreams. Mixing within the hepatocyte suspension and mass transfer of the dissolved gases will also be improved. The oxygen flux of silicone rubber is 500 GPUs (GPU=cm3/cm2-sec-cmHG x 10-6) a selectivity (O2 GPUs/N2 GPUs) of 2, while the CMS-7 membrane provides 9900 GPUs of oxygen flux and a selectivity of 2. Thus, use of the CMS-7 membrane dramatically increases oxygen flux. Figure 1. The Sybiol(R)Synthetic Bio-Liver Device --------------------------------------------------- | | | | | GRAPHIC | | | | Diagram of the Sybiol (R) Synthetic | | Bio-Liver Device | | | | | --------------------------------------------------- 5 3. Hepatocyte Function for Bioartificial Liver Support --------------------------------------------------- To date, the precise pathogenic mechanism of hepatic encephalopathy (HE) remains unknown. Observations in both animal models and patients with HE indicate that the syndrome is precipitated by the accumulation of toxins that are normally metabolized by the liver. High levels of these products promote neurotransmission failure. Various neuroactive substances, such as ammonia, (29) as well as endogenous or exogenous compounds with benzodiazepine agonistic properties (28) have been implicated as potential neurotoxins or agents, increasing the tone of the inhibitory gamma-aminobutyric acid neuro transmitter system. Drug metabolism is typically divided into two phases: Phase 1 oxidation or reduction reactions catalyzed primarily by P450 enzymes, and Phase 2 reactions which are conjugative reactions that increase aqueous solubility. In some cases, a drug may be sequentially metabolized by Phase 1 and Phase 2 reactions, as is the case with certain metabolites of diazepam. Other drugs such as acetaminophen are primarily metabolized by Phase 2 reactions (sulphation and glucuronidation). The primary requirement of cells in a liver support system is the preservation of the in vivo metabolic functions that prevent and/or decrease the progress of HE in patients with acute liver failure. Benzodiazepines, as well as other drugs or toxins, are metabolized in the liver by a family of enzymes that catalyze their oxidative degradation (Phase 1 metabolism) (31). This oxidation facilitates further detoxification steps (Phase 2 reactions) by which these byproducts are conjugated. The oxidative enzymes, known collectively as the P450 system, usually decay very rapidly in normal cultured hepatocytes (32). Most importantly, enhancement of diazepam (a benzodiazepine) and acetaminophen metabolism may support the utility of these BAL for liver support. Acetaminophen metabolism is higher in porcine hepatocytes than in human (Jauregui -unpublished data). In fact, acetaminophen toxicity is the etiology of acute hepatic failure most likely to be fully reversible by the provision of porcine hepatocyte based BAL support, obviating the need for transplantation (presented at American Association for the Study of Liver Disease Meeting, Nov. 1997). The study proposed here will explore the potential of an oxygen enhanced bioreactor culture system to support the metabolic function of porcine hepatocytes in vitro. We will address both Phase 1 and 2 reactions using 3 drugs for which the reaction products are well known and characterized, namely, diazepam, 7-ethoxycoumarin and actaminophen. The ability of porcine hepatocytes attached to microcarriers to convert ammonia to urea will also be assessed. 4. Oxygen Demands of Hepatocytes ----------------------------- The liver is one of the organs most dependent on oxygen. In fact, under basal metabolic conditions, 20-33% of the total oxygen utilized by the body is consumed by the liver (22). For this reason, hepatocytes and cell lines derived from hepatocytes (hepatoma) have high oxygen demands (10, 11). In vitro, the optimal conditions for cultivation and functioning of hepatocyte cell culture is 10-50% oxygen tension (10). Typically, mammalian cells consume oxygen at a rate of 0.05-0.5 (mu)mol/106 cells/hr (17). The reported oxygen uptake rate (OUR) for hepatocytes is approximately 1.0 (mu)mol/106 cells/hr, a high rate for mammalian cells (18). Oxygen must be continuously supplied because it is depleted quickly, even at relatively low cell densities. Oxygen is often the biolimiting nutrient in BALs, not just because of its high consumption rate, but also because, in typical cultures, it is found in relatively low concentrations. For example, the solubility of oxygen at 1 atm and 37(Degree)C is 0.2mM, while glucose is as high as 25mN (41). Oxygen requirements of hepatocytes vary depending on the culture phase, extracellular environment, cell density, cell type, and metabolic challenge. Low concentrations of oxygen in the early phase of hepatocyte cell culture establishment hinders cell attachment, spreading, and adaptation (8). Increasing the gas phase oxygen levels increases attachment and spreading; without sufficient oxygen, hepatocytes cannot assume the appropriate physical state for optimal albumin secretion (1, 8). The oxygen uptake rate (OUR) varies depending on the extracellular requirements; when hepatocytes were sandwiched between two collagen layers for long term testing, the OUR was 13.6 +/- 3 pmol/sec/(mu)g DNA in 2-16 hours (22). In BALs that contain hepatocytes within the hollow fiber module, high cell densities can cause a steep concentration gradient to form, so that cells far from sources of oxygen, nutrients, and metabolites may not receive sufficient amounts, causing, at best, loss of function, at worst, cell starvation and necrosis. Only the very first layer of cells can be assured sufficient oxygen (8, 11). In fact, when HF membranes are used for blood flow with cells cultured outside, oxygen is completely depleted at distances equal to one to two cell layers from the membrane outer wall (11). 6 Hepatocytes must have sufficient oxygen available to attach and spread and to perform metabolic functions, including specific liver functions that will benefit the patient. In the Sybiol(R) synthetic bio-liver device, cells are attached to microcarriers that recirculate through the hepatocyte loop. With the CMS oxygenator integrated into the hepatocyte loop, cells will regularly receive oxygen as the media and cells pass through the oxygenator. Since oxygen diffuses through the CMS membrane to dissolve in the liquid (media), all microcarriers should receive the same level of oxygenation. Thus, oxygenation will not be site dependent and steep concentration gradients can be avoided. The oxygenation technology we are proposing will allow control of the oxygen provided without a loss of operation over time due to wetting out of the oxygenator or fouling with biological matter, and in fact increases the oxygen flux. Another advantage our oxygenator provides is that this is a bubbleless source of oxygen. Sparging, which involves bubbling O2 directly into the media, often results in foam formation, and can damage cells (for details on the nature of the CMS-7 membrane, see below). The CMS oxygenation technology provides a "gentler" source of oxygen to cells removed from their natural environment. 5. Existing Oxygenation Techniques ------------------------------- The BALs that perfuse blood sometimes rely on red blood cells for oxygenation, although it has been suggested that additional oxygenation would improve hepatocyte function (12). In plasma-based models, red blood cells are usually excluded to reduce hemolysis, so an alternative oxygenation method is necessary. Many BALs incorporate either an indirect (in-series) or an integral (in-parallel) oxygenation system to maintain cell function and viability. With an indirect system, an oxygenator aerates either the plasma or medium with oxygen (sometimes pure) at pressures of 474-552 mmHg. High oxygen pressure, however, is problematic. Oxygen tension modulates enzyme activities, indicating that high O2 pressure could interfere with hepatocyte function. Since sparging can cause bubbles and foam formation that can damage the delicate hepatocyte cells, the preferred choice for oxygenation is a hollow fiber design, whether used as a separate, in-series unit (often a blood oxygenator) or integrated with a HF bioreactor. The materials employed for oxygenation in hepatocyte culture are typically hydrophobic polypropylene HF or silicone rubber (9, 14). These provide high oxygen flux with reduced bubbling and foam formation compared with spargers. However, there are adverse consequences of oxygenation with HF, including bubble formation in the cell mass due to changes in the pressure differences between the capillary spaces, and concentration of media due to evaporation of water at the air/media interface. In addition, the oxygen transfer is so high in the present oxygenation scheme that scavengers of free radical oxygenated species need to be added to the media. Coating the polypropylene fibers with CMS-7 would solve problems of bubble formation and media concentration while still permitting sufficient oxygenation and pH control. Furthermore, with CMS-7 treated fibers, as described in this proposal, hepatocyte and cell debris adhesion will be reduced, thereby minimizing fouling. This fouling problem is similar to the situation with blood oxygenators, the majority of which are based on non-wetting microporous polypropylene HF membranes. Blood flows on one side of the membrane, O2 on the other. Since the HF membranes are initially non-wetting, O2 and CO2 diffuse through the pores in the membrane. While these membranes function quite well initially, in many cases after 3-6 hours the holes in these porous HF become infiltrated with plasma. Pore blockage causes gas flux to drop significantly, plus plasma breakthrough into the gas phase occurs (15). We propose the use of a non-fouling membrane with higher oxygen flux in order to improve hepatocyte access to a bubbleless oxygen source, and, therefore, hepatocyte function and viability. Our CMS-7 coated HF membrane will provide a constant source of higher oxygen flux for a longer time period than those HF membranes currently in use. 7 6. New Integral Oxygenation Technology ----------------------------------- The cornerstone for developing an improved external oxygenation system using hollow fiber membranes (HF) in bioartificial liver assist devices (BALs) is a family of new perfluoropolymer membranes: perfluoro-2-2-dimethyl-1-3 dioxole (PDD) copolymerized with tetrafluoroethylene (TFE)(see Figure 2). These membranes will be employed in the HF oxygenator to enhance performance. Characteristics of these membranes include: 1) the ability to be made ultra-thin (0.5-2.0u) supported on flat sheet stock and/or porous hollow fibers, 2) chemical inertness and mechanical non-porous features which will prohibit cultured cells from attaching to the membrane, and 3) high oxygen and carbon dioxide permeability. HF treated with the PDD-TFE (CMS-7) polymer provide a bubbleless, non-fouling, biocompatible, rugged membrane that affords higher oxygen and carbon dioxide flux. a) SUPERIOR GAS-LIQUID TRANSFER- Through a number GRAPHIC of programs, we have demonstrated that the CMS-7 membrane provides oxygen transfer into liquid equal or superior to the performance of uncoated Diagrams of the modules (control modules without the CMS Structure of membrane). One of these projects involved using PDD-TFE the CMS membrane in bioreactors to enhance the O2 delivery and CO2 removal from fermentation broths. Table 1 summarizes the improvement (4x) of volumetric O2 mass transfer gained with the CMS membrane.
TABLE 1: Comparison of volumetric mass transfer coefficients for different contactors --------------------- ----------------------- ----------------------- ---------------- OTR Studies (k(l) a)* Coated Module (s(-1))** Uncoated Module (s(-1)) Spargers (s(-1)) --------------------- ----------------------- ----------------------- ---------------- Air 0.54 x 10(-3) 0.29 x 10(-3) 0.72 x 10(-3) Oxygen 3.3 x 10(-3) 0.88 x 10(-3) 5.5 x 10(-3) --------------------- ----------------------- ----------------------- ----------------
*K(1) a estimated from the time (sec.) for a volume of liquid (10L) to reach 63% air saturation measured with a dissolved oxygen probe. **Outside coated Sparging, another type of oxygenation, involves the physical dispersion of gas into liquid and, therefore, is more energy intensive. With a coated module, we approach the performance obtained by simple sparging. Sparging, however, is detrimental for a number of cell lines, since the bubbles cause unusually high shear stresses on the fragile cell walls, resulting in the breakage of the cell wall and cell death. We have also completed a project on oxygenation of water for aquaculture (DOC SBIR). With PVDF modules coated with CMS-7 on the fiber ID (inside diameter), O(2) flux rates for the coated device showed significant improvement compared to the uncoated devices, see Table 2. Additional work with Avecor polypropylene (PP) modules showed that the CMS membrane provided comparable oxygen mass transfer, see Table 3.
Table 2: Comparison of O2 flux rates for a PVDF coated with CMS-7 versus an uncoated module* -------------------------------------------------------------------------------------------- Test Description O(2) flux in gm/cm(2)/min %Increase Coated Uncoated in Flux Oxygenation with Pure O(2) (200 cm(2), 1000 um ID fiber) 0.28 x 10(-4) 0.18 x 10(-4) 56% Oxygenation with Air (200 cm(2), 1000 um ID fiber) 0.34 x 10(-4) 0.19 x 10(-4) 79% --------------------------------------------------------------------------------------------
*Runs made under identical operating conditions. 8
Table 3: Mass Transfer Testing of Polypropylene Modules with and without the CMS Membrane ---------------------------------------------------------------------------------------------------------- Module Membrane Membrane a Pressure Drop K(1) a (sec(-1)), average Number Status Thickness O(2)/N(2) 1 LPM 5 LPM 1 LPM 5 LPM ---------------------------------------------------------------------------------------------------------- 99-1191 - - - ~0 2.2 psi 2.0E-02 4.4E-02 99-1191 w/CMS-7 3.0 micron 2 ~0 2.3 psi 1.7E-02 4.6E-02 ---------------------------------------------------------------------------------------------------------- 99-1195 - - - ~0 2.0 psi 1.5E-02 4.6E-02 99-1195 w/CMS-7 2.3 micron 1.8 ~0 2.1 psi 1.4E-02 5.3E-02 ----------------------------------------------------------------------------------------------------------
Figure 3 schematically explains that this increase in gas transfer relates to the position of the gas-liquid interface. In an uncoated module, this interface exists at, or at a small distance into, the pore mouth. Molecules moving from the gas to the liquid phase must diffuse through this stagnant liquid phase into the edge of the liquid phase boundary layer, where they are carried away into the bulk of the solution by convection. In the case of the coated fiber, this stagnant liquid phase is replaced by the CMS polymer phase, which has considerably less diffusional resistance to gas molecules. As a result, the observed mass transfer rate is greater. The size of the diffusional boundary layer is also reduced, since the portion that would have been occupied by the liquid phase is now occupied by the CMS polymer. Figure 3: Gas Liquid Interface for CMS-7 Membrane on a Microporous Substrate GRAPHIC Diagram of the Gas Liquid Interface for CMS-7 Membrane on a Microporous Substrate b) COMPARISON TO OTHER COMMONLY USED HOLLOW FIBER MEMBRANES - In our studies, we focus on oxygen transfer rather than carbon dioxide transfer. Because our membrane's gas flux is so high, the transfer of oxygen into a cell culture medium becomes liquid-side limited. Therefore, if we optimize the transport of oxygen into the liquid, the faster permeating carbon dioxide will be maximized as well. Our CMS-7 polymer has nearly 40 times the oxygen flux of commonly used silicone rubber. We have also compared the oxygen transfer rate of our coated flat sheet membranes with microporous Gore-Tex(TM) and with existing non-porous silicone rubber materials under no-load conditions. The results are summarized in Table 4, and show a 2-3x improvement with CMS-7. Table 4 presents results of experiments with Sf-21 insect cells that have a high O(2) demand. The reported oxygen uptake rate (OUR) for Sf-21 cells is 0.24 (mu)mol/106 cells/hr compared to approximately 1.0 (mu)mol/10(6) cells/hr for hepatocytes. We observed a 3-4x higher cell density per unit membrane area for the CMS-7 membrane oxygenators versus the silicone rubber membrane oxygenator when used with bioreactors. We expect to see similar improvement with hepatocyte detoxification functions. Under non-bubble forming conditions, our non-porous CMS-7 membrane outperforms Gore-Tex(TM) with the measured O(2) delivery rate twice as much as Gore-Tex(TM) and three times as much as non-porous silicone rubber. We are able to successfully and reproducibly coat both the inside and outside of several different polymeric hollow fiber supports with CMS-7, as seen in Table 5. We have successfully demonstrated the feasibility of outside coating both hydrophobic (polypropylene) and hydrophilic (cellulose ester, polysulfone) HF by putting a thin layer (0.6-0.9u) of good selectivity (O(2)/N(2) ~ 1.7) on the outside of HF. There is no problem with fiber sticking with our coating procedures. 9
Table 4: Oxygen transfer rate for flatsheet materials/zero bubble formation ----------------------------------------- ---------------------- --------------------- --------------------- CMS-7 Gore-Tex(TM) Silicone Rubber (Non-porous) (Porous) (Non-porous) ----------------------------------------- ---------------------- --------------------- --------------------- Membrane Thickness (um) 1 ? 100 60 Gas Pressure (psig) 3 < 1 3 ml/L O(2) Delivery Rate ------*10(2) min 18.1 8.8 5.5 ----------------------------------------- ---------------------- --------------------- --------------------- Cell Densities of SF-21 obtained 2.1 x 10(4) 0.6 x 10(4) with Flat Membrane Sheets (Membrane area (Membrane area (cells/ml-cm(2) membrane area) =130 cm(2)) =260 cm(2)) ----------------------------------------- ---------------------- --------------------- ---------------------
Table 5. Outside -vs- Inside Coating of PDD-TFE on Various Polymeric Hollow Fiber Supports --------------- ------------- --------------------- -------------- ---------------- ----------- Unit Coating Support Area (cm(2)) Thickness(m) O(2)/N(2) =============== ============= ===================== ============== ================ =========== MCE-50-E Outside Cellulose Ester 50 0.7 1.7 MCE-50-D Inside Cellulose Ester 50 1.5 1.7 TexasLOD-A Outside Polysulfone 22 1.9 1.5 MPS-680-3 Inside Polysulfone 680 0.2 1.7 MPP-63-G19 Outside Polypropylene 63 0.8 1.8 MPP-250-21 Outside Polypropylene 250 0.9 1.7 MPP-250-G14 Outside Polypropylene 250 0.9 1.9 MPP-850-G23 Outside Polypropylene 850 0.6 1.7 MPP-1000-G1 Inside Polypropylene 1000 1.0 1.8 --------------- ------------- --------------------- -------------- ---------------- -----------
Table 6: Coating of Polypropylene Hollow Fiber Substrates from Various Manufacturers ---------------------------------------------------------------------------------------------------------- Supplier/ Surface Number Selectivity(O(2)/N(2)) Thickness (mm) Module Type Manufacturer Area(m(2)) Coated Average Std Dev Average Std Dev ---------------------------------------------------------------------------------------------------------- Terumo BLOX(a) Terumo 1.8 2 1.51 0.01 0.4 0.3 Liqui-Cel(b) Celgard 2.0 5 1.60 0.4 0.6 0.1 KrosFlow(1a) Spectrum, Inc. 1.0 5 1.69 0.06 0.4 0.2 KrosFlow(1b) Spectrum, Inc. 4.25 2 1.7 0.05 0.6 0.1 Avecor Affinity(2a) Avecor 2.5 8 1.6 0.05 1.5 0.5 Cellgas(1b) Spectrum, Inc. 0.25 5 1.66 0.1 0.3 0.2 IVOX(1b) OTD, Inc. 0.25 8 1.59 0.08 0.9 0.2 ----------------------------------------------------------------------------------------------------------
1 Fibers manufactured by Mitsubishi a Inside Coated 2 Fibers manufactured by Celgard b Outside Coated Table 5 discusses our ability to successfully inside or outside coat a wide range of PPs, suggesting that we have developed a universal process for coating PP. This table demonstrates the reproducibility of our success and the ability to scale it up to sizes consistent with large scale bioreactors. By contrast, others have been unsuccessful when coating the outside surface of microporous PP supports. Bob Schucker from Exxon reported at the November 1995 American Institute of Chemical Engineers annual meeting that they were unable to coat microporous PP with their polyurea/urethane membranes. Our success in coating CMS-7 onto microporous PP most likely relates to 1) the very low surface energy of the perfluoropolymer, 2) the excellent film forming capability of CMS-7, and 3) our superior fabrication process. Success in coating PP is especially valuable to this application since PP is inherently hydrophobic, a quality that will provide a "second layer of protection" should there be any defects in our coating. In addition, PP is tough and flexible, and it is already familiar to those in the bioreactor industry. 10 c) RESISTANCE TO FOULING AND DEPOSITION - The drop in mass transfer of microporous membrane-based gas-liquid contactors can be attributed to the wetting out of the substrate pores by the liquid phase (refer to Figure 3). Coating the membrane increases the time over which a contactor can be operated without having to be dried to recover the mass transfer performance of the original device. The device can now be operated with both phases at atmospheric pressure. In addition, one can now use these devices to carry out selective absorption of gases into liquids at high pressures, without requiring the pressurization of the gas phase to prevent the intrusion of the liquid phase into the pores of the microporous substrate. Compact Membrane Systems, Inc. has supplied modules to Cellex Biosciences, a company with the largest hollow fiber bioreactor market in the world. Cellex manufactures a gas-exchange cartridge (GEX) to reoxygenate the medium and remove CO2 for pH control. When Cellex uses a silicone sheet GEX, this accounts for 25-50% of the total manufacturing cost. Using a porous PP GEX is one quarter the cost, but this tends to wet out through the pores. Even in the absence of surface active foulants, the PP GEX wetted out after two days of use. Cellex is using a PP GEX coated with the CMS polymer in a bioreactor with a murine hybridoma (MH483) producing an IgM antibody, a passive anti-malarial immunization. The run is over 30 days old, and has produced 1.5g of antibody. The pH control was good, demonstrating sufficient CO2 exchange through the CMS GEX, and no loss in GEX efficiency was observed. No condensation was noted, and wet-out has not occurred. The CMS membrane significantly improved the long term performance of the GEX (20). d) RUGGEDNESS - CMS-7 also has excellent thermal stability, with a high Tg, above 200(degrees)C, so autoclave temperatures of 120(degree)C have little or no effect on it. Secondly, membranes tested before and after autoclaving show no significant change in gas flux performance in either direction. This indicates a very rugged material that can operate in a supported as well as unsupported mode. This ability to work in an unsupported mode is significant in gas/liquid applications, as gas transfer is higher if the liquid flows on the coated side of the membrane, allowing the gas to pass through an unsupported membrane. Because of our membrane's organophobic and hydrophobic properties, we tested its ability to operate in an oil/gas environment. We tested our nonporous CMS-7 and an enhanced microporous PTFE (e-PTFE) sheet. The gas flux of each membrane was measured and then they were submerged in a vacuum pump oil for at least two hours. The oil was then drained and the system was blown with air. The change in performance before and after oil coating for CMS-7 was modest; however, the flux of the e-PTFE dropped to zero when exposed to oil. We also observed breakthrough of the vacuum pump oil with the e-PTFE. This was due to oil getting into the pores of e-PTFE, hindering the flow of gas. By contrast, oil could not pass through the non-porous, organophobic, perfluoro CMS coating, and never reached the microporous support. The oil was easily washed from the surface of the CMS-7 coated membrane. Similar results were obtained using an air/oil aerosol. These tests demonstrate the ruggedness and fouling resistance of CMS-7. e) BIOCOMPATIBILITY TESTING - We have also conducted biocompatibility testing both in vitro and in vivo for the use of our membrane in extracorporeal membrane oxygenators (ECMO). Platelet adhesion was studied in vitro using a perfusion chamber to visualize fluorescently labeled platelet adhesion to CMS-7 coated glass coverslips. Whole blood from healthy, non-medicated donors was perfused over the slides, and non-adherent platelets were removed. Platelet deposition onto CMS-7 coated coverslips was not significantly different (p>0.05) than deposition onto uncoated glass coverslips (controls). To evaluate the thrombogenicity of the CMS-7 copolymer, coated and uncoated mini-oxygenators were perfused with radiolabeled platelets. Platelet deposition onto CMS-7 coated mini-oxygenators was less than that measured on uncoated (control) mini-oxygenators. However, this difference was not significant (p>0.05). 11 In vivo biocompatibility of CMS-7 coating used in ECMO was tested using calves. In this experiment, one coated and one uncoated (control) oxygenator were tested in parallel for 24 hours. The surgical procedure was well-tolerated and the animals recovered rapidly. All biochemical indices of organ function remained stable. Figures 4 and 5 summarize CO2 transfer results for the coated and uncoated membrane oxygenators. Similar findings were obtained as regards VO2. However, because the inlet pCO2 for the data shown below is in close conformity with AAMI standards for venous CO2 content during oxygenator testing, VCO2 data are presented here for discussion. The trend regarding CO2 transfer in uncoated, HF oxygenators follows the usual, well-documented pattern of a dramatic loss in gas exchange capacity over a period of hours (15). On the other hand, VCO2 in the CMS-7 coated membrane oxygenator is virtually unchanged at 20 hours versus the very early perfusion time points. Further, VCO2 for the coated units at these same very early times matches the CO2 transfer rates for the control. This latter finding supports the in vitro data above regarding minimal loss of mass transfer capability in HF covered with CMS-7. Further, the VCO2 data suggest that the perfluoropolymer may resist the fluid infiltration and plasma breakthrough associated with currently available hollow fiber membrane oxygenators, at least for the duration of in vivo perfusions approaching 24 hours. Figure 4: Figure 5: [GRAPHIC] [GRAPHIC] Diagram of the Uncoated MO Diagram of the Perfluoropolymer In Vivo Perfusion Time (Hours) Coated MO In Vivo Perfusion Time (Hours) We believe the improved oxygen flux delivered by a CMS-7 coated HF will benefit hepatocytes used in BALs. Oxygenation will be bubbleless, and should not damage cells. The hepatocytes in BALs are especially fragile because they have been removed from their native environment. Any "protection" that can possibly be provided to these cells will encourage greater viability and function. The coated membrane is non-fouling and resists liquid penetrations, so it should outperform the commonly used polypropylene and silicone rubber, where fouling and wet out can be problematic. Its ruggedness assures the ability to withstand repeated sterilization for long term use. For these reasons, CMS-7 coated hollow fiber membranes should improve the access of hepatocyte cells to oxygen, thereby improving their viability and function, and the modules should remain operational for longer periods than oxygenators containing uncoated hollow fibers. 7. Technical Objectives -------------------- The overall goals of this program are to develop and test CMS-7 membrane modules for O2 and CO2 transport for use as an in-series oxygenator in a BAL hepatocyte loop. We will demonstrate enhanced hepatocyte function compared to use of oxygenators without the CMS-7 membrane. The key technical goals of the Phase I program are: 1) Design and fabricate CMS oxygenators using polypropylene CMS-7 membranes. 12 2) Demonstrate that the CMS oxygenator permits sufficient microcarrier passage. 3) Demonstrate that the CMS oxygenators have the ability to transfer oxygen/carbon dioxide into/from water and hepatocyte media at a rate equal or superior to that of oxyenators without the CMS-7 membrane (controls). Determine which type(s) and configuration(s) are most appropriate for use with the Sybiol(R) Bio-Liver Device. 4) Demonstrate that the CMS-7 coated membrane has a broader operating range (greater resistance to bubble formation or wet out [liquid penetrations]) than uncoated polypropylene membranes. 5) Demonstrate short term performance of the CMS oxyenator assessed by viability of a recirculating hepatocyte suspension. 6) Demonstrate a significant increase in primary porcine hepatocyte function with the CMS oxygenator by using the CMS oxygenator and the Sybiol(R) Bio-Liver Device as compared to performance of control oxygenators. In Phase II, we will: a) Conduct extensive evaluations of hepatocyte function. Long term (ten days) performance will be a key goal. We look forward to working with MultiCell Associates, Inc. (MCA) and Exten Industries in this area. b) Consider using a human hepatocyte cell line developed by MCA and/or use of a medium with higher O2 carrying capacity for Phase II. c) The Sybiol(R) Bio-Liver Device will be evaluated in vivo with use of the CMS oxygenator for provision of liver support function in a well-documented large animal model. d) Food and Drug Administration (FDA) approval will be secured for a Phase I clinical trial for device safety. This will set the stage for Phase III of the development, which will consist of Phase I and II FDA clinical trial assessment of the device for safety, dose, and efficacy. 8. Significance of Phase I Effort ------------------------------ This Phase I effort is composed of three work stages: 1) Fabrication and performance evaluation of the CMS oxygenator, 2) Enhancement of hepatocyte functional viability when the CMS oxygenator is used in conjunction with the Sybiol(R) Bio-Liver Device, and 3) Assessment of program results and planning for Phase II. More specifically, Phase I will demonstrate our ability to fabricate CMS oxygenators by applying the CMS-7 membrane to commercially available polypropylene hollow fiber modules. We will determine which type(s) of modules and membrane configuration(s) are most appropriate for use with the Sybiol(R) Bio-Liver Device based on initial gas testing, microcarrier passage, and oxygen flux into water and cell culture media. Furthermore, we will demonstrate that performance of these CMS oxygenators is equal or superior to comparable oxygenators without the CMS membrane (controls) in oxygen flux/mass transfer. Phase I will also demonstrate that these HF with the CMS membrane provide a broader operating range (greater resistance to bubble formation and liquid wet-out) than uncoated HF PP. In addition, HF with the CMS-7 membrane will perform for a longer duration than fibers without the membrane because CMS-7 is non-fouling, thus providing prolonged function. Thus, use of the CMS oxygenator will enhance performance of hepatocytes in the Sybiol(R) Bio-Liver Device. Prolonged BAL performance will provide the treatment the patient with acute or chronic hepatic failure requires. C. RELEVANT EXPERIENCE Dr. Nina Lamba (Principal Investigator) - Through her graduate and post-doctoral studies, Dr. Lamba developed an extensive knowledge of bioengineering, and the advancing field of tissue engineering. Recently appointed as the Director of the Biomedical Research Division of Compact Membrane Systems, Inc., she is overseeing projects to develop oxygenation devices for extracorporeal and intravenous application and an artificial liver project based on a dual fiber bioreactor. These programs are also funded through the NIH SBIR program. The information gathered from these investigations can be directly applied to advance the development of membrane devices to oxygenate bioartificial organs. 13 Dr. Stuart M. Nemser (President - Compact Membrane Systems) - Dr. Nemser has 25 years experience in the development and commercialization of membrane technologies. Dr. Nemser identified and patented the family of PDD polymers for separations. Related to PDD-TFE copolymer development, E.I. DuPont holds key patents on the PDD monomer and associated polymers. Dr. Nemser has a strong, close working relationship with the key DuPont groups and DuPont is committed to exclusively supplying Dr. Nemser with PDD-TFE copolymers for evaluation in key membrane separation applications. In his current position as President of CMS, Dr. Nemser is focused on commercializing this family of PDD polymers. Compact Membrane Systems, Inc. - CMS is currently involved in two key research areas that relate directly to this program: 1) PDD-TFE copolymer development and 2) thin film membrane development. Markets where CMS is presently active include membranes to facilitate oxygen transfer to blood and to an artificial liver, and membranes to facilitate ozone to water for disinfection of food processing plants and dental equipment. We anticipate starting a BAL project on developing a dual fiber bioreactor (integral O2 verus the in-series oxygenation proposed here) shortly. Compact Membrane Systems, Inc. has also made excellent commercialization progress since receiving its first Phase II SBIR in the 4th quarter of 1994. Our accomplishments during the last five years include the following: (1) Sales/royalties have doubled each of the last four years and exceeded $290,000 in 1999. (2) We have established a long term licensing and supplier agreement with Pall Corporation (sales of $1,000,000,000). This agreement will be supplying between $100,000 and $200,000 minimum annual royalty for their sales into the semiconductor market using our chemically resistant bubbleless gas delivery system. Equally important, Pall Corporation has agreed to supply us with membrane systems. This now provides us with a large quality supplier to drive our technology into the semiconductor and other markets. (3) We have established a strong relationship with Praxair and their membrane subsidiary Innovative Membrane Systems, Inc. (IMS). Praxair is a global leader in industrial gases ($5,000,000,000/yr) and they provide us with excellent market access and needed module manufacture. CMS and IMS in collaboration with a major industrial diesel manufacturer have invested over $150,000 developing an improved CMS membrane. (4) We are commercially supplying Celgard/Hoechst standard degassing modules. Sales to date to Celgard/Hoechst have been in excess of $100,000. They have introduced our technology as a separate product line for degassing low surface energy fluids. (5) Our portable membrane systems for supplying oxygen for respiratory care has been developed and will be submitted shortly for FDA approval. Five separate companies have presented proposals to CMS for joint commercialization of this technology. One company (Chad Therapeutics) commissioned a market research study that identified a market opportunity of $30,000,000 (20,000 units per year at $1500 per system). (6) E.I. DuPont, from whom we have licensed the original technology, continues to show interest in our technology. DuPont has provided CMS in excess of $100,000 of key equipment (e.g. GC and mass spectrometer) and access to other facilities (e.g. SEMs and computer models). They also have supplied us with research materials that are not available commercially or developmentally. (7) We have filed nine patents and four have been either issued or allowed. These patents will provide us with a sustainable competitive advantage. 14 Please refer to the Biographical Sketches and Bibliography section for details on key personnel, including consultants. D. EXPERIMENTAL DESIGN AND METHODS The protocols described below will demonstrate the improved performance achieved with the Sybiol(R) system by improving the function of the in-series oxygenator. It is critical for CMS and its collaborators to demonstrate such advantages before moving to commercialization. 1. Fabrication of CMS Oxygenators ------------------------------ Compact Membrane Systems, Inc. (CMS) will fabricate improved oxygenators appropriate for use with the Sybiol(R) device. Subtasks include: 1. SELECT MODULES APPROPRIATE FOR USE WITH MICROCARRIERS - The primary design consideration will be choosing a configuration which will allow passage of the hepatocyte microcarriers. Although Phase I hepatocyte testing will not use microcarriers, CMS plans to design the oxygenator so microcarrier use is possible. Exten Industries has used microcarriers in the past, but has found that in a recirculating system, attachment is not necessary for function (42). For future commercialization, CMS believes that the ability to function with or without microcarriers, for use with either Sybiol(R) or similar devices, will be advantageous. Typical microcarriers such as those proposed for the Sybiol(R) device are 150-200(mu)m in diameter. This requires the development of an oxygenator which will not entrap microcarriers. There are at least two approaches to avoid this problem. First, we could use modules with a lower packing density (fewer fibers per unit area) so microcarriers could flow freely on the outside (shell) of the fibers with oxygen/air on the inside (lumen). The second design option is to use large diameter fibers so that the perfusate with microcarriers can pass through the lumen with gas in the shell. The former is preferable to obtain higher mass transfer rates. CMS has a number of module options available to address this issue. We plan to purchase commercially available polypropylene (PP) hollow fiber (HF) modules from Celgard (see attached letter). In a routine process, CMS currently applies the CMS membrane to Celgard PP modules for commercial sale by Celgard for industrial liquid degassing. However, for this BAL project we will require modules with either a lower fiber packing density or a larger fiber inside diameter (ID). Although CMS does not anticipate problems with availability of Celgard modules in the appropriate configurations, other options are available to us should the Celgard modules not be feasible. We have experience with in-house fabrication of modules using commercially available fibers, and have fabricated such modules in a variety of configurations, sizes, and with a variety of different fiber materials and IDs. In addition, recently CMS acquired Intravenous Oxygenator (IVOX) fabrication equipment and will soon begin in-house fabrication for a Phase II NIH SBIR project. This equipment and the knowledge gained from this project will be applicable to in-house module fabrication. This would permit design and fabrication of either low packing density or large fiber diameter modules. CMS also has access to polysulfone fibers in larger IDs with the CMS membrane applied to the outside from Innovative Membrane Systems, Inc (IMS)(see attached letter). Once the CMS membrane is applied to the polysulfone support, "wet out" is no longer a problem. Thus, CMS does not anticipate problems with obtaining the appropriate module configurations for oxygenator fabrication. 2. APPLY THE CMS MEMBRANE - Applying the CMS-7 membrane to HF modules is a routine procedure at CMS. First the CMS-7 polymer is dissolved in the appropriate perfluorosolvent, then pushed through the module shell (for outside coating) or fiber bore (for inside coating). The solution is then removed from the module, taking into consideration the need to remove all residual solvent and polymer. After allowing the module to sit for an appropriate period of time, the module is dried to remove any residual solvent. Our in-situ technique for applying the CMS membrane lends itself to easy scale up and multiplicity of production units. As shown in Tables 5 and 6, CMS is able to coat either the inside or outside of a variety of module types. Key variables that we will consider are coating solution concentration and in-situ coating conditions, such as solution pressure and flow rate. We will focus on assuring that there is no significant change in fiber ID while developing high transmembrane gas transport. We will also be sensitized to any issues associated with biocompatibility and potential system leaks. 15 3. TEST GAS/GAS PERFORMANCE OF MODULES - After these membranes have been fabricated, we will do initial non-destructive, in-house, gas-gas testing to determine if these modules are meeting our target performance goals. We routinely use oxygen/nitrogen selectivity and membrane thickness (calculated based on nitrogen GPUs, Gas Permeation Units) to establish membrane quality. 4. TEST PERFORMANCE WITH MICROCARRIERS - Successfully fabricated modules will be tested for performance with unseeded microcarriers to visualize microcarrier flow patterns. Using a dye to color the microcarriers will make tracking microcarrier flow easier. This subtask will allow us to determine which modules will perform best with the microcarriers. If any problems are noted, we will consider other module sources and configurations. The best performing module(s) will be used for mass transfer testing in Task 2. 2. Testing of Oxygen Flux of CMS Oxygenators ----------------------------------------- Modules successfully fabricated in Task 1 will be used for initial mass transfer testing to determine oxygen flux and operating range. With the current silicone rubber-based passive oxygenation system used for in Sybiol(R) device, control of oxygen flux is not possible. For each module type or configuration a comparable module without the CMS membrane will be used as a control. Initial oxygen flux testing will use an oxygen sweep on the unsupported (no CMS-7 membrane) side of the PP fibers and water on the supported/membrane side. The rate of oxygen uptake with time and the mass transfer coefficient (K1a) will be measured. In our studies, we focus on oxygen transfer rather than carbon dioxide transfer, because if O2 transport is optimized, the faster permeating CO2 will be maximized as well. The pressure will be kept low enough to prevent bubbling. Water and media flow rate will be similar to the flow rate currently used for the Sybiol(R) system, and we will vary oxygen flow to determine optimal flow rate and pressure. In testing conducted for an aquaculture project, CMS observed that oxygen removal from water did not differ significantly between modules with and without the CMS membrane, but that differences could be observed between manufacturers. We are currently developing oxygen mass transfer testing as a means to determine membrane suitability for gas/liquid applications, so it is appropriate to include such testing. After preliminary tests with water, we will repeat these tests using Chee's essential media, the cell culture media we will use for the hepatocytes. To test the operating range, the gas pressure will be increased and a suitable surface wetting agent will be added. Increased pressurization tends to cause fouling and wet out (liquid penetrations) with uncoated PP. The experiment will then be rerun as described above. With this task, we will demonstrate which module(s) provides the most appropriate oxygen flux into water and media, and what conditions (ie oxygen flow rate and pressure) are needed. 3. Hepatocyte Isolation -------------------- Primary porcine hepatocytes will be isolated by established techniques (27). Briefly, a modification of Seglen's two step collagenase digestion procedure will be used. Female Yorkshire swine (~12 kg) will be anesthetized, intubated, and supported with supplemental oxygen. The liver perfusion will be performed in situ under sterile conditions. The liver will be perfused via the vena cava with a calcium free buffer until clear of blood. Collagenase buffer solution will be perfused until the liver is visibly digested and then excised. 16 The liver will be minced gently and washed three times with an iced buffered salt solution and sedimented by centrifugation. The final hepatocyte slurry will be resuspended in supplemented Chee's essential media (Gibco, GrandIsland, NY) to a concentration of ~35 x 10(6) cells/ml. Typical hepatocyte yield is 17 x 10(9) at 85% viability. Porcine hepatocytes have been chosen for these investigations based on the following: 1. MultiCell Associates (MCA) has nearly 10 years of experience with porcine hepatocytes having developed in vitro growth and maintenance requirements and developed proprietary procedures for their isolation, culture and cryopreservation. 2. Porcine hepatocytes are the cell of choice in most ongoing clinical trials of temporary liver support systems. Vitagen, Inc. uses their proprietary human tumor-derived C3A cell line, which in a study comparing them to porcine hepatocytes in vitro expressed much lower detoxification functions. 3. Human hepatocytes, while theoretically optimal, are in limited supply. The scarcity of donor human livers for transplantation severely limits the potential to procure human hepatocytes in commercially feasible numbers for temporary liver support systems. In addition, the function of human hepatocytes is highly variable. Specifically, enzymatic function decays substantially within the first 24-48 hours after cell isolation. Unlike porcine hepatocytes, the methodologies to procure and freeze large quantities of human hepatocytes have not been well established. MCA has tested human hepatocyte cultures produced by Biowhittaker/Clonetics, S. Strom and others but found them to be of poor quality, with no P450 detoxification expression in vitro. Currently, we feel they are a poor choice of cell to use for liver assist device design and optimization. Rat hepatocytes would be more readily available and far less costly than pig, but in our experience and that of others they can not be easily scaled-up to clinical size devices, are not representative of either human or porcine hepatocyte function; and have very different requirements for in vitro maintenance than either pig or human hepatocytes. MCA is developing an immortalized nontumorigenic human hepatocyte cell line which may be used in future studies, however, their stage of characterization makes it premature to use. The zoonotic concerns raised by the use of porcine hepatocytes remain unresolved. The findings that porcine endogenous retrovirus (PERV) (Patience, et al., 1997) could infect human cells in vitro led to a moratorium on clinical trials involving porcine tissue. The FDA is now examining each clinical trial on a case-by-case basis and has allowed many to go forward including the Phase III pivotal of Circe Biomedical's porcine hepatocyte-based extracorporeal liver support system, "HepatASSIST." Among the requirements for these studies to continue are the archiving donor and recipient sera and tissues and patient follow-up protocols. While France and the United Kingdom are still on clinical holds, clinical trials have been reactivated in Holland, Germany, Italy, and Belgium. While the debate continues, we must proceed cautiously under the prevailing guidelines and regulations of the FDA, but we should not rule out the clinical use of porcine-derived hepatocytes at this time. 4. In Vitro CMS Oxygenator Performance with Recirculating Hepatocytes ------------------------------------------------------------------ The Sybiol(R) Bio-Liver Device, as described in section B.2. and diagrammed in Figure 1, is based on recirculation of hepatocyte cells in suspension. Preliminary tests will include a recirculating loop of primary porcine hepatocytes isolated as described above with oxygen delivery provided in-series by the CMS oxygenator. This will allow for initial cell viability testing before integrating with the entire Sybiol(R) Bio-Liver Device to test detoxification function. In all experiments, monloayer cultures will be established as controls, maintained and tested in parallel to the bioreactor cultures to ascertain viability of the isolated hepatocyte population. To reduce the number of animals needed for this study, only the most promising one or two CMS oxygenators will be used for this testing. We may also choose to use smaller scale/lower volume systems for these test. In these studies, the freshly isolated cells will be inoculated into Chee's modified media at a density of 35 x 10(6) hepatocytes/cc of system volume. Initial viability will be assessed with Trypan Blue staining. Viability of the hepatocytes will be monitored over the culture period by lactate dehydrogenase (LDH) leakage into the media reservoir.(35) Assay of LDH leakage offers a non-invasive method to evaluate plasma membrane integrity confirming that metabolism of the test substrate is not due to "persistence of microsomal activity in non-viable cells." 17 Morphologic evaluation of the hepatocytes maintained will also be performed by sampling the cell culture media at regular intervals. Samples will be extracted and the cells evaluated using fluorescein diacetate and propidium iodide labeling to detect viability by fluorescence microscopy. We recognize the critical importance of the multiple in vitro environmental factors which can influence cell function beyond simple oxygenation. In fact, in an effort supported by Circe Biomedical to develop the biological component of their liver support system.(36,37); MCA scientists have studied media formulations, substrate requirements, and other conditions of porcine hepatocyte culture to develop proprietary methods for their isolation, culture, and cryopreservation (USPN5,795,711; USSn 08/5,411,462). A key objective of this program is to use our high gas flux non-porous membranes to supply needed dissolved oxygen in order to maintain hepatocyte viability. If cell viability is not maintained by the testing procedure described above we will first consider exploring other types of modules and membrane configurations available to address this problem. If these modifications are not successful we will consider media related modifications. Previous studies using rat hepatocytes showed Chee's media supported pO2 levels up to 100 mmHg. We will soon begin a similar NIH SBIR project on bioartificial livers (based on a dual fiber bioreactor) that should address issues of oxygen carrying capacity of the media; thus, for this project, we will have established the media limitations and requirements. Nevertheless, if higher O2 levels are necessary for porcine hepatocytes, we will explore the incorporation of hemoglobin and/or perfluorinated species to improve oxygen carrying capacity. (38,39) 5. In Vitro CMS Oxygentator Performance with the Sybiol(R) System -------------------------------------------------------------- The CMS oxygenator will be placed in line with the complete Sybiol(R) Bio-Liver Device (see Fig. 1), including the HF cartridge that acts as a cell culture media/blood interface. Before using the device for patients, trial runs are recommended to test the hepatocyte suspension quality. Standard operating procedures for these trials will be the model for this task.(40) The Sybiol(R) will be filled with 250ml of hepatocyte suspension in the hepatocyte loop, with 300ml of albumin solution or frozen plasma in the blood loop, functioning in a "closed loop" configuration. A stock solution of 100 mM ammonium chloride will be prepared, and 5.5ml will be added to the blood loop. Urea production will be measured. Three additional substrates will be used: acetaminophen, diazepam and 7-ethoxycoumarin (7-EC). (Refer to the "Significance" section for details on hepatocyte function). Levels of albumin (a hepatocyte specific protein) in culture media will also be measured by ELISA. Diazepam (50 ug/ml), 7-EC (50 mg/mL), and acetaminophen (5mM) will be added to the blood loop. These preliminary tests will last 6-8 hours. Prior testing with the Sybiol(R) device has demonstrated sufficient urea production (83% of ammonium chloride added) after six hours (40). The cultures will be incubated with the test substrates for three hours, at which time media samples will be collected from each and stored at -30(degree)C for later assay. Diazepam and 7-EC metabolites will be analyzed by high performance liquid chromatography (HPLC) methods described previously by Jauregui et al (33). Acetaminophen will also be determined by HPLC (34). The conversion of ammonia to urea will be measured by a kit assay (Sigma, St Louis, MO). As with Task 4 we will monitor hepatocyte viability over the culture period by lactate dehydrogenase (LDH) leakage into the media reservoir. Monolayer cultures will be established as controls to ascertain functional integrity of the isolated hepatocyte population. Performance of the oxygenation fibers (based on ability to deliver the required oxygen, resistance to breakthrough, etc.) will also be monitored during testing. If possible, once testing is complete, the oxygenators will be returned to CMS for oxygen transfer rate testing to see if any changes have occurred while the oxygenators were in use. Results will be statistically analyzed using ANOVA and student's t test. 18 6. Evaluation of Results and Preparation for Phase II -------------------------------------------------- Success will be determined by our ability to fabricate polypropylene hollow fiber modules appropriate for oxygenation of the Sybiol(R) Bio-Liver Device that do not severely inhibit microcarrier passage and that provide oxygen transfer equal or superior to that of comparable modules without the CMS-7 membrane. We will demonstrate that, when the Sybiol(R) system is used in conjunction with the CMS oxygenator, hepatocyte viability is comparable or superior to that of cells oxygenated with control oxygenators. We will also demonstrate that the CMS oxygenator significantly enhances primary porcine hepatocyte phase 1 and 2 detoxifications when the CMS oxygenator is placed in-series in the Sybiol(R) system. Furthermore, we will demonstrate that the CMS oxygenator provides longer term performance due to resistance to wet out and fouling. This will set the stage for Phase II, which will include more comprehensive short-term testing of primary and immortalized hepatocytes (MCA has developed a cell line that functions best in suspension), longer term functionality tests, and in vivo animal trials. Working with MCA and Exten Industries will be important factors in determining project success.
Timeline of Scheduled Tasks and Responsibilities: ----------------------------------------------------------------------------------------------------------- Tasks Months Key 1 2 3 4 5 6 7 8 9 10 11 12 Responsibility ----------------------------------------------------------------------------------------------------------- 1) Fabricate Oxygenators ----------> CMS 2) Oxygen Mass Transfer Testing -------> CMS 3) Hepatocyte Isolation --------> MCA 4) Test with recirculating hepatocytes ----------------> MCA 5) Test with Sybiol(R) ----------> MCA/Exten 6) Plan Phase II ----> CMS/MCA/Exten -----------------------------------------------------------------------------------------------------------
F. VERTEBRATE ANIMALS MCA's (Dr. Jauregui) primary facility (~ 4500 sq ft.) is located in Warwick, RI where activities focus on the development of human origin cell lines for therapeutic and diagnostic applications. Additionally MCA has a contractual agreement with Rhode Island Hospital (RIH) leasing ~750 sq ft of office and laboratory space for their porcine hepatocyte-based activities. Through this agreement, MCA has access to RIH's AALAC accredited Central Animal Facility which includes sterilization, surgical, and husbandry facilities. Additionally, the RIH-MCA agreement provides for MCA access to numerous Core Research Laboratory services including: electron microscopy, digital imaging, flow cytometry etc. on a fee for service basis. MCA anticipates IACUC review and approval of the animal procedures described in this Phase 1 SBIR proposal to be completed by May 2000. G. CONSULTANTS - See attached letters H. CONTRACTUAL AGREEMENTS Contractual agreements have been established with Dr. Hugo Jauregui of MultiCell Associates, Inc. and Barbara Corbett of Exten Industries. I. LITERATURE CITED 1. Bonn, D. Hybrid devices offer hope for the failing liver. The Lancet. Nov 16, 1996. 348(9038): 1372(1). 2. Rotem, A, M Toner, S Bhatia, BD Foy, RG Tompkins, ML Yarmush. 1994. Oxygen is a factor determining in vitro tissue assembly: Effects on attachment and spreading of hepatocytes. Biotechnol Bioeng. 43:654-660. 3. Sussman, NL, JH Kelly. 1996. Artificial liver support. Clin Invest Med. 19(5):393-99. 19 4. Jauregui, HO, C.J-P Mullon, BA Solomon. 1997. Extracorporeal artificial liver support. Chapter 30. In Principles of Tissue Engineering. RP Lanza, et al. R. G. Landes Co. Austin, TX: 463-479. 5. Qiang, S, Y Yaotiong, L Hongyin, H Klinkmann. 1997. Comparative evaluation of different membranes for the construction of an artificial liver support system. Int J Artif Org. 20(2):119-124. 6. Gerlach, JC, N Schnoy, J Vienken, M Smith, P Neuhaus. 1996. Comparison of hollow fibre membranes for hepatocyte immobilisation in bioreactors. Int J Artif Org. 19(10):610-616. 7. Gerlach, JC. 1996. Development of a hybrid liver support system: a review. Int J Artif Org 19(11):645-654. 8. Catapano, G. 1996. Mass transfer limitations to the performance of membrane bioartificial liver support devices. Int J Artif Org 19(1):18-35. 9. Flendrig, LM, AA TeVelde, RAFM Chamuleau. 1997. Semipermeable hollow fiber membranes in hepatocyte bioreactors: A prerequisite for a successful bioreactor? Artif Org. 21(11): 1177-1181. 10. Nishikawa, M, J Uchino, M Michiaki, M Takahashi, K Taguchi, M Koike, H Kamachi, H Kon. 1996. Optimal oxygen tension conditions for functioning cultured hepatocytes in vitro. Artif Org. 20(2):169-77. 11. Smith, MD, DE Cairns, R Veitch, RB Cousins, JDS Gaylor. 1997. Analysis of oxygen transfer in hollow fibre hepatocyte bioreactors. Artif Org. 21(6): 531(219). 12. Gerlach, J, K Kloppel, P Stoll, J Vienken, C Muller. 1990. Gas supply across membranes in bioreactors for hepatocyte culture. Artif Org. 14(5):328-333. 13. Smith, MD, D Cairns, JM Courtney, RB Cousins, MH Ekevall, MH Grant, JDS Gaylor. 1997. Characterisation of a hybrid artificial liver bioreactor with integral membrane oxygenation. Artif Org. 21(6): 531 (220). 14. Gerlach, JC, J Encke, O Hole, C Muller, JM Courtney, P Neuhaus. 1994. Hepatocyte culture between three dimensionally arranged biomatrix-coated independent artificial capillary systems and sinusoidal endothelial cell co-culture compartments. Int J Artif Org. 17(5):301-306. 15. Cheng, BT, EF Leonard. 1995. Light microscopic visualization of plasma intrusion into microporous HF's. ASAIO J. 41:863-872. 16. Kasai, S, M Sawa, M Mito. 1994. Is the biological artificial liver clinically applicable? A historic review of bioartificial lvier support systems. Artif Org. 18(5):348-354. 17. Miller, WM, HW Blanch. Regulation of animal cell metabolism in bioreactors. In Animal Cell Bioreactors. Eds. CS Ho, DIC Wag. Stoneham, MA: Cutterworth-Heinemann, 1991. 127. 18. Peng, HA, BO Palsson. 1996. Determination of specific oxygen uptake rates in human hematopoietic cultures and implications for bioreactor design. Ann Biomed Engin. 24:373-381. 19. Smith, MD, AD Smirthwaite, DE Cairns, RB Cousins, JD Gaylor. Jan 1996. Techniques for measurement of oxygen consumption rates of hepatocytes during attachment and post-attachment. Int J Artif Org. 19:1, 36-44. 20. Graham, M. 1998. Testing a CMS-coated gas-exchange cartridge in a hollow fiber system. Cellex Biosciences. Personal communication. 21. Ash, SR. Aug 1998. Hemocleanse, Inc. Personal communication. 22. Rotem, A, M Toner, RG Tompkins, ML Yarmush. 1992. Oxygen uptake rates in cultured rat hepatocytes. Biotech and Bioengin. 40:1286-1291. 23. Algenix, Inc. 1998. Company Profile. [http://wsh2.cems.umn.edu/Algenix2/ page2.html] 24. Majumdar, S, KK Sirkar. 1992. Hollow-fiber contained liquid membrane. In Membrane Handbook. WSW Ho, KK Sirkar. Chapman & Hall, NY. 25. Jauregui, HO, S Naik, H Santaggini, J Pan, D Trenkler, C Mullon. 1994. Primary cultures of rate hepatocytes in hollow fiber chambers. In Vitro Cell Dev Biol. 30A:23-29. 20 26. Liu J, P Jing, S Naik, H Santangini, D Trenkler, N Thompson, A Rifai, JR Chowdhury, HO Jauregui. Characterization and evaluation of detoxification functions of a Nontumorigenic Immortalized Hepatocyte Cell Line (HepLiu). Submitted to Cell Transplantation, 1998. 27. S Naik. 1996. Isolation and culture of porcine hepatocytes for artificial liver support. Cell Trans. 5(1) 107-115. 28. Basile, AS, EA Jones, P Skolnick. 1991. The pathogenesis and treatment of hepatic encephalopathy: Evidence for the involvement of benzodiazepine receptor ligands. Pharmacol. Rev. 43:27-71. 29. Butterworth, RF, JE Gigu6re, J Michaud, J Lavoie, GP Layrargues. 1987. Ammonia: Key factor in the pathogenesis of hepatic encephalopathy. Neurochem. Pathol. 6:1-12. 30. Clayton, DF, M Weiss, JED Darnell. 1985. Liver-specific RNA metabolism in hepatoma cells: Variations in transcription rates and MRNA levels. Mol. Cell. Biol. 5:2633-2641. 31. Kaplowitz, N. Drug metabolism and hepatotoxicity. In: Kaplowitz, N., ed. Liver and Biliary Diseases. Baltimore: Williams & Wilkins; 1992:82-97. 32. Reid, LM, DM Jefferson. 1984. Culturing hepatocytes and other differentiated cells. Hepatology. 4:548-559. 33. Jauregui, HO, SF Ng, KL Gann, DJ Waxman. 1991. Xenobiotic induction of P-450 PB-4 (11B1) and P-450c (1A1) and associated monooxygenase activities in primary cultures of adult rat hepatocytes. Xenobiotica. 21:1091-1106. 34. Colin, P, G Sirois. 1986. Simultaneous determination of the major metabolites of styrene and acetominophen, and of unchanged acetominophen in urine by ion-pairing high-performance liquid chromotography. J. Chromatogr. 377:243-251. 35. Jauregui, H.O., Hayner, N.T., Driscoll, J.L., Williams-Holland, R., Lipsky, M.H., Galletti, P.M. 1981. Trypan blue dye uptake and lactate dehydrogenase in adult rat hepatocytes - freshly isolated cells, cell suspensions, and primary monolayer cultures. In Vitro 17)12):1100-1110. 36. Jauregui, H.O., Naik, S., Santangini, H.A., Trenkler, D.M., Mullon, C.J.-P. The use of microcarrier-roller bottle culture for large-scale production of porcine hepatocytes. Tissue Engineering 3(1):17-25; 1997. 37. Naik, S., Santangini, H.A., Trenkler, D.M., Mullon, C.J.-P., Solomon, B.A., Pan, J., Jauregui, H.O. 1997. Functional recovery of porcine hepatocytes after hypothermic or cryogenic preservation for liver support systems. Cell Transplantation 6(5):447-454. 38. Lowe, KC; Davey MR, Powell JB 1998 "Perfluorochemicals: Their applications and benefits to cell culture; Trends in Biotechnical; June (6) 272-7 39. Smith, MD, Smirthwaite, Ad, Cairns DE, Cousins RB, Gaylor JD 1996. "Techniques for measurement of oxygen consumtion rates of hepatocytes during attachment and post attachment; J. of Artificial Organs, 19/1: 36-44. 40. Teca SOP No. 96-01 for Sybiol. Section 4, pg. 4-10. 41. Macdonald, JM, Griffin, JP, Kubota, H, Griffith, L, Fair, J, and Reid, LM. "Chapter 21: Bioartificial Livers." 252-286. 42. Bhatia, SN, MD, PhD. March 2000. Assistant Professor of Bioengineering, Division of Gastroenterology/Hepatology, University of California, San Diego. Personal communication. 21