A hollow-fiber bioreactor is a 3-dimensional cell-culturing system based on hollow fibers, which are small, semi-permeable capillary membranes arranged in a parallel array with a typical molecular weight cut-off (MWCO) range of 10–30 kDa. These hollow-fiber membranes are often bundled and housed within tubular polycarbonate shells to create hollow-fiber bioreactor cartridges. Within the cartridges, which are also fitted with inlet and outlet ports, are two compartments: the intracapillary (IC) space within the hollow fibers, and the extracapillary (EC) space surrounding the hollow fibers. Cells are seeded into the EC space of the hollow fiber bioreactor and expand there. Cell-culture medium is pumped through the IC space and delivers oxygen and nutrients to the cells via hollow-fiber membrane perfusion. As the cells expand, their waste products and CO2 also perfuse the hollow fiber membranes and are carried away by the pumping of medium through the IC space. As waste products build up due to increased cell mass, the rate of medium flow can also be increased so that cell growth is not inhibited by waste-product toxicity. Because thousands of hollow fibers may be packed into a single hollow-fiber bioreactor, they increase the surface area of the cartridge considerably. As a result, cells can fill up the EC space to densities greater than 108 cells/ml. However, the cartridge itself takes up a very small volume (oftentimes the volume of a 12-ounce soda can). The fact that hollow-fiber bioreactors are very small and yet enable incredibly high cell densities has led to their development for both research and commercial applications, including monoclonal-antibody and influenza-vaccine production. Likewise, because hollow-fiber bioreactors use up significantly less medium and growth factors than traditional cell-culture methods such as stirred-tank bioreactors, they offer significant cost savings. Finally, hollow-fiber bioreactors are sold as single-use disposables, resulting in significant time savings for laboratory staff and technicians.
History In 1972, the Richard Knazek group at the NIH reported how mouse fibroblasts cultured on 1.5 cm3 hollow-fiber capillary membranes composed of cellulose acetate were able to form 1-millimeter-wide nodules in 28 days. The group recorded the final cell number as approximately 1.7×107 cells from a starter batch of only 200000 cells. When the same group cultured human choriocarcinoma cells on polymeric and silicone polycarbonate capillary membranes totaling less than 3 cm3 in volume, the cells expanded to an amount approximating 2.17×108 cells. The Knazek group was awarded the patent for hollow-fiber bioreactor technology in 1974. Based on this patented technology, companies began building different and larger-scale hollow-fiber bioreactors, with significant development and technological improvement occurring in the late 1980s to early 1990s. By 1990, at least three companies were reported to offer commercially available hollow-fiber bioreactors. One engineering advance included adding a gas-exchange cartridge, which enabled better control of system's pH and oxygen levels. Similar to a mammalian lung, the gas-exchange cartridge efficiently oxygenated the culture medium, allowing the bioreactor to support higher numbers of cells. Combined with the ability to add or remove CO2 for precise pH control, the limitations commonly associated with large-scale cell culture were eliminated, resulting in densely packed cell cultures that could be maintained for several months. In addition, control of the fluid dynamics within each hollow-fiber bioreactor led to further optimization of the cell-culture environment. By alternating the pressure gradient across the hollow-fiber membrane, media could flow back and forth between the EC side (cell compartment) and the IC side (hollow-fiber lumen). This process, combined with the axial media flow created when media passes down the length of the fibers, optimized the growth environment throughout the entire bioreactor. This concept is termed EC cycling, and was developed as a solution to the gradients that form within hollow-fiber bioreactors when media is pushed down the lengths of their fibers. Higher hydrostatic pressure at the axial end (media entering the fiber lumen) compared to the distal end of the bioreactor creates a Starling flow in the EC space, which is similar to what is observed in the body. This phenomenon also creates a nutrient-rich axial region and a nutrient-depleted distal region within the bioreactor. By incorporating EC cycling, the effects of Starling flow are eliminated, and the entire bioreactor becomes nutrient-rich and optimized for cell growth. Optimal IC and EC space perfusion rates must be achieved to efficiently deliver media nutrients and growth supplements, respectively, and to collect supernatant. During the cell-growth phase within these bioreactors, the media feed rate is increased to accommodate the expanding cell population. More specifically, the IC media perfusion rate is increased to provide additional glucose and oxygen to the cells while continually removing metabolic wastes such as lactic acid. When the cell space is completely filled with cells, the media feed rate plateaus, resulting in constant glucose consumption, oxygen uptake, and lactate production rates.
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