Tools and Applications of Biochemical Engineering Science

ferent strategies have been proposed in the literature to circumvent the problems of filter plugging. Avgerinos et al. [8] employed adherent cell lines growing on microcarriers to get aggregates (200–600 mm) that were far larger in size than the mesh opening employed (120 mm). By doing this, they were able to operate the filter for at least 31 days with no fouling problems. Varecka and Scheirer [86] used wire cages with 53 mm mesh opening and cells of approximately 20 mm size, and observed no sign of fouling for more than two months of continuous operation.

Esclade et al. [87] evaluated the influence of the screen material on fouling of spin-filters. They compared, under the same culture conditions, screens made out of stainless steel and polyamide with similar pore sizes (10 and 11 mm, respectively), similar screen surfaces and same porosity.After measuring the number of living and dead cells, and the protein and nucleic acids content in the deposits found on the screens after one week of recirculation experiments (recycle of perfusate without addition of fresh medium), they verified that the polyamide screen was superior to the stainless steel one, because it presented a low binding of proteins, nucleic acids and cells (either alive or dead). Scanning electron micrographs revealed that in the stainless steel screen nearly all the pores were filled with veils and cells, whereas in the polyamide screen the pores were practically free of cells and veils. They related this different behaviour to the different surface charge densities of both materials, metals being highly positively charged, polyamide being more neutral and cells rather negatively charged.

Analogous results concerning screen material were obtained by Avgerinos et al. [8]. In different experiments using stainless steel screens with various pore sizes (44–105 mm), the spin-filter became plugged after only 11 to 21 days of operation, with cells attached and growing on both sides of the mesh. Contrastingly, five perfusion cultivations up to 54 days in duration were carried out with a polymeric ethylene-tetrafluroethylene (ETFE) screen (120 mm pore size) without filter fouling. The authors concluded that the hydrophobic nature of the ETFE screen apparently minimized the attachment of cells to the screen and thus greatly extended the life of the spin-filter.

Beyond screen material and pore size, operating conditions have significant effects on filter clogging. Deo et al. [13] showed that fouling is heavily influenced by spin-filter rotation. They observed that the maximum perfusion rate achievable without clogging correlated with the square of the tangential velocity at the screen surface. Esclade et al. [87] observed that high perfusion rates increased screen clogging by increasing the drag towards the screen surface. They found in recirculation experiments (without fresh medium addition) that at a perfusion rate of 2 d–1 filter plugging just started after a viability decrease due to medium depletion occurred. Contrastingly, at a perfusion rate of 7 d–1 plugging took place instantaneously due to sucking of the cells onto the screen. Similar results had already been obtained by Tolbert et al. [88], who verified that at high medium withdrawal rates there was a buildup of subcellular debris in the range of the filter pore size, limiting the maximum perfusion rate and thus the maximum viable cell concentration attainable.

Fouling caused by increasing perfusion rates can be partially compensated by increasing the spin-filter rotation speed. However, according to Yabannavar

et al. [89], too high rotation speeds may increase cell leakage and, thus, there is an optimal intermediate rotation speed, which assures effective cell retention and unclogged filter operation.

Besides fouling, cell retention efficiency is another important factor regarding spin-filter performance. A poor retention capacity of a spin-filter leads to a low apparent growth rate of the cell culture due to cell leakage (Eq. 10).

For a spin-filter based bioreactor, an overall cell balance as given by Eq. (11) can be applied [89]:

where b, the cell passage factor, is the ratio between cell concentration in the perfusate and in the bioreactor, or (1 – E), according to Eq. (6).

Based on Eqs. (10) and (11), it is possible to relate the apparent cell growth rate to the perfusion rate and the retention efficiency E (Eqs. 12 and 13):

Iding et al. [90] observed an apparent cell growth rate of 0.27 d–1 after perfusion started, as compared to m = 0.65 d–1 in the batch phase. Contrastingly, Avgerinos et al. [8] took advantage of the aggregating propensity of CHO cells, which formed clumps 200–600 mm in size, to operate with a high separation efficiency using a 120 mm pore-size filter. Under these conditions, the specific rate (0.026 d–1) at which cells were lost through the screen was negligible in comparison to the apparent specific growth rate of the culture (0.26 d–1).

Yabannavar et al. [89] employed Eq. (11) to calculate the maximum perfusion flow rate attainable before cell wash-out, which occurs when cell growth is compensated by cell passage through the spin-filter or, according to Eq. (12), when mapp = 0. Under these conditions:

As a consequence, the perfusion bioreactor can only be operated up to a cell concentration supported by the perfusion rate Dmax . In this way, spin-filter retention efficiency determines the maximum attainable cell concentration in a given perfusion process.

Scale-up of spin-filters has been often studied in the recent years, and simple scale-up procedures have been proposed by Yabannavar et al. [81] and Deo et al. [13]. Yabannavar et al. [81] suggested a procedure based on keeping the ratio of permeation drag to lift force constant for the different production scales, in

order to avoid fouling. In cross-flow filtration [84], this ratio is defined as in Eq. (15):

Yabannavar et al. [81] proposed a proportionality relationship valid for spin-fil- ters based on an analogy to Eq. (15). They defined the Reynolds number based on the tangential velocity at the screen surface. Since in spin-filters the permeation velocity, or perfusion flux, is given by Eq. (16), and it can be assumed that the screen porosity e will be maintained constant throughout the scale-up process, it is possible to write a proportionality relationship for the ratio from drag to lift force in spin-filters as given by Eq. (17).

Eq. (18) can be obtained by expressing the variables in Eq. (17) in terms of geometric dimensions:

As bioreactor scale-up is usually accomplished by applying geometric similarity

and assuming fixed aspect /ratios (dsf /dbioreactor , hsf /hbioreactor , dbioreactor /hbioreactor),

Eq. (18) reduces to:

or, in terms of bioreactor volume:

Yabannavar et al. [81] used Eqs. (19) and (20) for scale-up purposes. Based on successful operation conditions determined for an existing 12-L bioreactor, they calculated the spin-filter dimensions and operation conditions for an existing 175-L bioreactor. Experiments with the spin-filter designed for the large-scale bioreactor resulted in an absence of filter clogging with cell retention efficiency similar to the 12-L bioreactor. This was considered as an evidence that the suggested scale-up strategy is adequate.

Deo et al. [13] proposed an empirical model based on experimental results obtained with a small-scale spin-filter based bioreactor. They observed that the perfusion flux capacity, which was defined as the perfusion flux at which fouling begun to occur, was proportional to the inverse of the cell concentration and to the square of the tangential velocity at the screen surface (Eq. 21):

Since the perfusion flux is given by Eq. (16) and the screen porosity can be assumed to be constant, the proportionality given in Eq. (22) is obtained.

V X

Eq. (22) shows that, at fixed tangential velocity and cell concentration, perfusion capacity increases proportionally with the A/V ratio. Alternatively, maintaining the tangential velocity and the cell concentration constant, an increasing bioreactor volume requires a proportional increase in the spin-filter area in order to maintain the perfusion rate D constant. Deo et al. [13] employed this strategy to scale-up a process to the 500-L scale on the basis of the operating conditions established in a 7-L bioreactor. They considered the strategy successful, since the scaled-up process showed the same temporal profiles of cell growth rate, maximum cell concentration and product level as in the small scale.

Although much progress has been made in the last decade regarding operation, design and scale-up of spin-filters, in most works found in the literature either fouling or retention problems (or both) were observed. A better comprehension of the fluid and particle dynamics involved in spin-filter perfusion would improve this situation. In this context, a valuable tool that could be used is computational fluid dynamics (CFD), which has been recently employed for the design and performance prediction of other cell separation devices [46,114].

6

Tangential Flow Filtration

In tangential flow filtration (TFF) systems, feed solution flows perpendicular to the permeation direction. The tangential flow generates high shear rates at the filter surface, contributing to avoid filter fouling and thus enabling a continuous operation and a high filtrate flux. TFF units usually consist of hollow fiber (HF) or flat plate (FP) membrane systems.

Since the early work by Radlett [91], many works have been published on the use of membrane-based TFF for both harvesting [1, 92, 93] and perfusion of mammalian cells (Table 6). According to these works, membrane-based TFF presents various advantages over other separation devices:

–process simplicity and easy adaptation to currently available stirred-tank bioreactors [94];

–high filtration capacity with complete cell retention, generating a particlefree harvest and thus facilitating further downstream processing [1];

–external operation, which enables aseptic filter replacement if fouling occurs, and represents no constraint to process scale-up, such as in the case of internal retention devices [95];

–low operating costs and ability to use the same process for a large amount of different products without great additional development work [92].

However, in spite of the cleaning action generated by the tangential flow, shear rates and thus feed flow rates are limited by the shear sensitivity of the cell line employed. Thus, fouling remains a major concern when developing a TFF-ba- sed perfusion process. Although different authors reported no significant cell damage during TFF processes with either FP or HF units [1, 92, 96], limitations in flow rates due to adverse shearing effects have been frequently reported.

Maiorella et al. [93] carried out tests with flat plate and hollow fiber units with various cell lines (NS-1 myeloma cells, three different hybridoma cell lines and SF-9 insect cells). They observed in all cases that cells were not damaged in laminar flow at average wall shear rates up to 3000 s–1. However, cell damage started to occur beyond this shear level. Zhang et al. [94], using a hollow-fiber cartridge for a hybridoma perfusion culture, observed that shear damaging effects on cell growth occurred for a recirculation rate of 700 mL/min, which corresponded to a calculated maximum shear rate at the fiber walls of 1266 s–1.

Because shear rates and feed flow rates are limited by the shear stability of cells, other strategies have been proposed to avoid fouling. Kawahara et al. [97] reduced the filtrate flux by increasing the membrane area, and adapted a washing line to their perfusion system. In order to wash accumulated cells and debris out of the system, a daily washing process was carried out at a flow rate 80 times higher than that used to pump the cell suspension. Even though, an increasing membrane clogging was observed from the 20th day on and culture had to be terminated on the 30th day.

Smith et al. [98] employed three different techniques to reduce fouling, and suggested that the third one was the most effective method to avoid blocking of the fibers:

–backflush on the HF cartridge shell side;

–increased recirculation rate through the fiber lumen;

–pulsing the feed, whereby care was required to prevent rupture of the fibers when using this technique.

Zhang et al. [94] also implemented different methods to reduce membrane blockage and to prolong cultivation time:

–membranes were coated with polyethylene glycol, which reduces protein adsorption and cell adhesion;

–periodical backflushing of the HF cartridge was carried out;

–a cell-bleeding line was used during backflushing to remove cell debris accumulated in the cartidge.

Combining these techniques, they carried out cultivations for 250–350 h, and were able to repeatedly use the same cartridge (four times at least) without measurable deterioration in filtration efficiency. However, when perfusion rate and cell concentration in the bioreactor increased, fouling eventually occurred. Van Reis et al. [92] provided backpressure on the filtrate line to control filtrate rates and so to avoid too high initial filtration rates, which can cause rapid fouling. De la Broise et al. [99] compared the filter performance using membranes of different pore sizes (2 and 10 mm). In both cases partial retention of the produced IgM was observed and membranes had to be changed every 5 days, the

fouling effects being more accentuated when using the 2 mm membranes. Even though, it was possible to carry out cultivations for about 30 days.

Maiorella et al. [93] observed that fouling behavior was dependent on cell size. For smaller cells or suspensions containing significant levels of debris, high flux rates could not be maintained without inducing high transmembrane-pres- sures. This behavior is in qualitative agreement with the hydrodynamic lift theory, since lift velocity is predicted to increase with particle diameter to the second [100] or third power [101].

Scale-up of TFF processes is usually done by arranging a large number of short devices in parallel. The individual filters should be kept relatively small in axial dimension, helping to avoid the excessive build-up of feed pressure resulting from the pumping of viscous slurries [96]. By arranging TFF membranes in parallel or in series, units can be scaled-up to process thousands of liters of cell culture fluid in a few hours [1].

Van Reis et al. [92] reported the scale-up of a HF system for the recovery of human tissue plasminogen activator (t-PA) produced by recombinant CHO cells from the 2.5-m2 to the 180-m2 scale.A robust and reproducible process was achieved by combining linear scale-up principles, control of fluid dynamic parameters and experimentally defined limits of product retention, which meant maintaining channel length, wall shear rate and flux constant.

Although the fouling problem seems not yet to be completely solved, the simple scale-up, combined with reproducibility of repeated TFF runs, membrane reusability, and suitability of a TFF system for different cell lines and culture media [1] make TFF an attractive choice for mammalian cell separation.

7

Dynamic Filtration

As discussed previously, minimization of filter clogging and fouling requires the generation of sufficient wall shear stress at the filter surface. However, cell lysis has to be avoided in mammalian cell cultivation, in order to allow high cell densities and viabilities to be obtained. Therefore, optimization of animal cell filtration requires careful adjustment of the wall shear stress at the whole membrane area [104]. Dynamic filtration is based on creating a relative motion between the membrane and its housing or between the membrane and a rotor, so that the shear rate produced by this relative motion is independent of feed flow rate [105]. Basically, there are dynamic filtration devices of two different geometries: vortex flow filters (VFF) and rotating disc filters (RDF). A third filtration device which takes advantage of the relative motion between membrane and cell suspension is the double membrane stirrer (DMS) developed by Büntemeyer et al. [106]. Although this is not considered a traditional dynamic filter, it will be briefly discussed in this item. An overview of reports that used dynamic filters for mammalian cell perfusion is given in Table 7.

VFF devices consist of a stationary cylinder, inside which a concentric cylinder rotates. The rotating movement of the inner surface of the annular gap creates a Taylor-Couette flow [16], generating Taylor vortices. The filter medium can

158 L.R. Castilho · R.A. Medronho

be mounted either on the rotating or, more commonly, on the stationary cylinder. The stable Taylor vortices generated induce mixing perpendicular to the membrane surface and reduce the extent of concentration polarization and fouling. The flow rate passing through the annulus can be kept low, because the fouling preventive action comes from the Taylor vortices and is thus not coupled to the feedstream flow [16]. VFF have also been called external or ex-situ spin-fil- ters [20, 107].

RDF units consist of a rotating disc inside a stationary housing, the membranes being mounted either on the rotating disc [108, 109] or on the housing [104, 110]. Such rotating disc devices have been shown to be very effective not only for solute recovery from microbial suspensions [111, 112], but also for mammalian cell separation [104, 113]. Like in VFF, this concept permits the independent adjustment of wall shear stress, transmembrane pressure and residence time, allowing a straightforward optimization of the filtration process and the establishment of a controlled shear field. Therefore, it has also been called CSF – controlled shear filtration [104].

The core of double membrane stirrer perfusion bioreactors is a stirrer on which two microporous hollow fiber membranes are mounted, one of them being hydrophobic and used for bubble-free aeration, the second of them being hydrophilic and used for cell-free medium exchange [15]. This system has been reported to provide viable cell densities of 20 million cells per milliliter for more than two months [106]. Although Lehmann et al. [15] have described the scale-up of this system to the 20-L and 150-L scale, it has been most commonly employed at the bench-scale.

Rebsamen et al. [138, 139] were among the first to investigate the use of vortex flow filters for mammalian cell separation. They carried out experiments with suspensions of hybridoma and human breast carcinoma (SKBR-5) cells. For rotor angular velocities up to 500 rpm no viability decrease was observed. A concentration of SKBR-5 cells under sterile conditions with subsequent cultivation of the concentrated cells showed that the cells were able to grow normally after being subjected to vortex flow filtration. The authors suggested that the VFF could be employed as a long-operating-life cell retention device in perfusion cultures or, in a miniaturized version, as an on-line sampling device.

Hawrylik et al. [140] investigated the influence of rotor speed, retentate flow rate and system back pressure on the separation of U-937 human lymphoma cells expressing interleukin-4 (IL-4) on the cell surface, aiming at a reduction of shear damage to cells and a minimization of contaminating cellular proteins in the process stream. They verified that even under the gentlest conditions applied no evidence of clogging or buildup of differential pressure across the membrane occurred. Under these conditions, the cell surface IL-4 binding capacity was completely recovered and a minimal loss of cell viability was observed. These facts suggested that the cells remained intact during filtration. The authors concluded that VFF proved to be an efficient, low shear method for mammalian cell separation and medium harvesting.

Mercille et al. [141] compared TFF and VFF for the perfusion cultivation of a hybridoma cell line producing an IgM, either using serum-supplemented or protein-free medium, with and without the addition of DNAse. They observed