Journal of Controlled Release 64 (2000) 81–90www.elsevier.com/ locate / jconrel

Matrices for tissue engineering-scaffold structure for abioartificial liver support systema , a,b b a*J. Mayer , E. Karamuk , T. Akaike , E. Wintermantel

aChair of Biocompatible Material Science and Engineering, Wagistrasse 23, CH 9852 Schlieren, ETH Zurich, SwitzerlandbFaculty of Bioscience and Biotechnology, TIT, Tokyo, Japan

Received 1 April 1998; accepted 11 May 1999

Abstract

This study proposes a new composite scaffold system. A woven polyethylenterephtalate (PET) fabric was coated on oneside with a biodegradable PLGA film, in order to obtain a geometrically polarized scaffold structure for an bioartificial liversupport system. The composite structure ensures the stability of the membrane during degradation of the membrane polymer.The mesh size of the composite does not significantly influence the degradation behavior. Hepatocyte culturing studies revealthat the formation of aggregates depends on the mesh size and on the pretreatment: The largest aggregates could be observedafter 48 h when PVLA coating, large mesh size and EGF were combined. Thus, the combination of a geometricallystructured, partially degradable scaffold with receptor-mediated cell attachment sites offers promising possibilities in livertissue engineering. 2000 Elsevier Science B.V. All rights reserved.

Keywords: Tissue engineering; Degradables; Hepatocytes

1. Introduction allow one to design an arrangement of architecturalelements, such as pores and fibers. Therefore, vari-

Biocompatibility of materials has been defined as ous textiles arranged into superstructures have beencompatibility of a technical with a biological system proposed for tissue engineering scaffolds [2,3] be-[1], in the quest for structural compatibility as well cause they have the ability to provide optimal spatialas for surface compatibility of the implant. Applied and nutritional conditions for cell maintenance.to tissue engineering, avital scaffold structures are Because of the shortage of donors, there is a needdesigned to take up cells and to provide them with for techniques that provide both long term and shortnutrients. Scaffolds should mimic the tissue architec- term liver support to prevent acute liver failure. Anture as well as the extracellular matrix (ECM) effective artificial liver support system should pro-components of the target tissue. Textile materials vide the essential processes of the liver, e.g. syn-

thetic and metabolic functions, detoxification andexcretion [4,5]. Dialyzer-like artificial liver systems,*Corresponding author. Tel.: 141-1-633-6311; fax: 141-1-633-which replace only this single liver function, lack the1120.

E-mail address: mayer@biocomp.mat.ethz.ch (J. Mayer) metabolic and synthetic functionality of the liver

0168-3659/00/$ – see front matter 2000 Elsevier Science B.V. All rights reserved.PI I : S0168-3659( 99 )00136-4

82 J. Mayer et al. / Journal of Controlled Release 64 (2000) 81 –90

[6–8] and, thus, the survival rate of patients with have been proposed by many authors using differentfulminate hepatitis does not exceed 30% [4,8]. pore-forming mechanisms, such as salt leaching [18–Consequently, the use of isolated hepatocytes is 20] and solvent foaming [21]. Degradation rates ofenvisaged to replace a broader spectrum of liver- PLGA polymers depend on many physical andspecific metabolic functions. Currently, implantable chemical factors and can be adjusted from daysas well as extracorporeal, so called bioartificial liver (amorphous, low-molecular-weight) to months (crys-(BAL) support systems are considered. In both talline, high-molecular-weight). Since mechanicalapproaches, isolated liver cells are used in a variety properties are not of major concern in the currentof configurations, e.g. suspensions [5,9], attached to application, amorphous polymers with relatively lowsubstrates [10–12] as well as encapsulated in micro- molecular weights were selected (see Table 2).sphere- [4,13] or hollow fiber systems [14–16]. Furthermore, degradable polymer membranes are

Compared to implantable systems, BAL systems considered to serve as delivery systems. The mem-allow better control of the cellular environment, brane can be modified by receptor-specific ECMparticularly with respect to transport properties. In coatings e.g. glycopolymers. Applied to polystyreneaddition, immunological rejection is expected to be culture dishes, these glycopolymers have been shownreduced because of the potential to separate the to influence the attachment and differentiation ofrecipient’s lymphocytes from the exogenous hepat- primary hepatocytes [22,23]. Investigations usingocytes by plasmapheresis. A potential disadvantage uncoated nylon fabrics [24] and collagen-sandwichesof an extracorporeal system is the need for vascular [25] as scaffolds indicated an interaction between theaccess, including the risk of thromboembolic compli- scaffold architecture and the cell culture. Therefore,cations. A central task in the development of BAL it is suggested that the foreseen three-dimensionalsystems is the appropriate design of the cell carrier structure of the proposed scaffold will influence thesystem, which should have a high density of fully aggregation behavior and, thereby, the functionalitydifferentiated hepatocytes while providing optimal of hepatocytes.nutritional conditions. The aim of this paper is to introduce the concept

In this communication, a composite scaffold is of a composite scaffold system as a support structureproposed as a carrier system for the liver cells. The for a bioartificial liver assistance device. The paper isscaffold consists of a non-degradable, woven poly- focused on characterization of the scaffold propertiesmer fabric that is characterized by a well-defined and provides only preliminary qualitative data re-mesh size. The fabric is coated on one side with a garding the performance of the composite scaffold inbiodegradable polymer matrix membrane. Thin primary hepatocyte cultures.membranes have promising characteristics becausethe degradation rate, permeability due to micropor-osity and the surface chemistry can be controlled by 2. Materials and methodsthe fabrication. Therefore, a composite of non-de-gradable meshes with a degradable membrane results Woven 131 fabrics (polyethylenterephtalate, PET;in a minimized membrane thickness while mechani- Sefar AG, Switzerland) were selected as the non-cal stability during degradation is optimized. degradable part of the scaffold, and had mesh sizes

Biodegradable PLA/PGA polymers show a that varied between 1000 and 20 mm (Table 1). Thebiphasic degradation behavior with a fast degrading fabrics were cut into circular pieces with a diameterbulk and a slower degrading surface. This behavior of 80 mm. After sonication in ethanol, the fabricswas shown to be strongly size-dependent and a were dried at 568C for 3 h and residual stresses werecritical thickness of 300 mm was estimated for pure cured at 1008C. Two PLGA copolymers and onesurface degradation [17]. Therefore, thin (,100 mm) PDLA homopolymer (Boehringer Ingelheim, Ger-and highly porous devices such as membranes are many) were chosen to cover a wide range ofsupposed to show a uniform degradation behavior different degradation rates (Table 2). The coatingbecause of minimized matrix volume. Porous bio- process was performed as follows: Sugar was dis-degradable polymeric scaffolds for tissue engineering solved in distilled water at room temperature in a 1:1

J. Mayer et al. / Journal of Controlled Release 64 (2000) 81 –90 83

Table 1Properties of the woven PET fabrics that were used to prepare composite scaffolds

Mesh size Open area Wire Weight Thickness2(mm) (%) diameter (mm) (mg/mm ) (mm)

50 33 38 0.037 60200 43 104 0.072 170720 45 350 0.150 745

Table 2Selected properties of the three amorphous PLA/PGA polymers

Polymers Molecular Glass Estimatedweight (g /mol) transition degradation time

Poly(D,L-lactide) |200,000 50–658C 24 weeks to 18RESOMER R207 months

Poly(D,L-lactide–co-glycolide), 85:15 |200,000 45–608C 90–220 daysRESOMER RG858

Poly(D,L-lactide–co-glycolide), 50:50 |20,000 40–558C 7–60 daysRESOMER RG503

(w/w) ratio, with stirring. After sonication, the sugar were measured by gel permeation chromatographysolution was cast into glass petri dishes and the (GPC, Knauer, Germany; Column, Polymer Lab-fabrics were placed onto the film. The polymer oratories Plgel 5 mm Mixed-C; length, 600 mm;solutions (55 mg/ml in CHCl , 10 ml) were cast inner diameter, 7.5 mm; particle size, 5 mm; in-3

onto the fabrics. The solvent was removed by drying jection volume, 100 ml; pressure, 52 bar; flow-rate, 1the samples for 12 h under vacuum at room tempera- ml /min). PLA/PGA polymers were diluted to 0.2ture and, subsequently, for 12 h at 508C. The vol.% in tetrahydrofuran (THF, Fluka). Poly(styrene)

6contribution of the meniscus to the film area was (Polymer Laboratories; M , 580–1.28 10 g/mol;w

estimated from scanning electron microscope (SEM) polydispersity, 1.03) was used as the standard.cross-sectional images. b-Galactose carrying styrene homopolymer,

For evaluation of film degradation, square pieces, PVLA (Fig. 1), was added to aqueous solutions at215315 mm , were placed in 12-well plates. Degra- concentrations of 100, 10 and 1 mg/ml, as described

dation medium (phosphate-buffered saline; PBS)was added in a w/v ratio of approximately 25 mgfilm/ml buffer. Thus, the surface /volume ratio was

24.5 cm /ml buffer, as required in ISO-10993-9 forin-vitro degradation experiments. Degradation con-ditions were set at pH 7.4, 378C and 5% CO . The2

medium was replaced every four days. After selecteddegradation times, samples were taken out, washedwith distilled water and subsequently lyophilized for24 h before being subjected to the various analyses.

Weight loss was determined from the lyophilizedspecimens and was calculated relative to the initialdry weight. Water adsorption was measured on wetspecimens after gently wiping the residual waterfrom the surface. Fig. 1. Chemical structure of PVLA, a b-galactose-carrying

Molecular weight (M and M ) and polydispersity styrene homopolymer.w n

84 J. Mayer et al. / Journal of Controlled Release 64 (2000) 81 –90

in Ref. [22]. A 100-mg/ml stock solution was determined to be around |95%, as measured byprepared and concentrations were made by diluting. trypan blue exclusion. Cells were seeded on coatedUsing the quartz microbalance technique [26], it was meshes in serum-free Williams Eagle mediumnot possible to quantitatively record the adsorption of (Gibco) either with or without epidermal growthPVLA coating because water absorption by the factor (EGF; 50 ng/ml, EGF was kindly donated bypolymer film during coating prevented equilibration Hitachi Chemicals, Japan). The cell suspension was

5of the system. adjusted to 1.5?10 cells /ml and to a seeding density4 2Scaffold discs (Ø, 15 mm) were dipped in ethanol, of 5?10 cells /cm . Culture plates were incubated for

vacuum dried and sterilized using UV light (6 h on 4 h to induce initial cell attachment. After theeach side). The samples were fixed onto a 24 well- medium was aspirated, 500 ml of fresh mediumplate using the tip of a forceps that had been covered containing EGF (50 ng/ml) were added. Mediumin sterilized (autoclaved) silicone grease. The scaf- was replaced daily. Cell morphology was character-folds were washed twice with 500 ml of sterilized ized by means of optical and scanning electronPBS (-). Sterilized (0.45 mm membrane filtration) microscopy. No quantitative, biochemical tests wereaqueous solutions (1 and 100 mg/ml; 500 ml /well) performed at this preliminary stage of the inves-of PVLA were added to the wells after being tigation.sonicated for 10 min. The films were coated for 2 hat room temperature and were subsequently washedwith 500 ml of PBS. Bovine serum albumin (BSA, 3. ResultsSigma, 0.1% in PBS) was added at 378C for 30 min,to prevent non-specific cell attachment. Having The composite scaffolds exhibit a polarized struc-removed the BSA solution, the samples were washed ture, with a plain ‘film side’ and a highly organizedtwice with PBS. Primary hepatocytes were isolated ‘fabric side’, with well-defined cavities of betweenfrom ICR mice in situ according to the two-step 800 and 50 mm, depending on the mesh size of thecollagenase procedure of Seglen [27]. Viability was fabric (Fig. 2). The thickness of the cast films was

Fig. 2. SEM micrographs of the coating of woven PET fabrics with biodegradable polymers. Different mesh widths [720mm (a) and 50 mm(b)] and polymers [RG503 (a) and R207 (b)] are shown. The film thickness is around 80 mm and the fabrics are coated on one side only.

J. Mayer et al. / Journal of Controlled Release 64 (2000) 81 –90 85

homogeneous inside the mesh. Approaching thethread, the formation of a meniscus could be ob-served, however, the contribution of the meniscus tothe film decreased with increasing mesh size (Table3). The wall height of the formed cavities is relatedto the fabric architecture and, thereby, varied be-tween 50 and 150% of the respective thread diameter(Table 1). The anchorage of the films in the PETfabrics was found to be purely mechanical. Table 3also shows the relationship between film thicknessand the specific mass gain for the individual meshes.

The films appear transparent after casting anddrying. No major difference in weight loss (Fig. 3)was detected for different fabric geometries. PDLLA(R207) and 85:15 PLGA (RG858) polymers behavesimilarly when comparing a given mesh size. Afteran initial weight loss of 2% within two days, the ratedecreased to less then 1% in ten days. The 50:50PLGA (RG503) scaffold showed a similar behaviorfor ten days. However, a rapid weight loss wassubsequently observed, which resulted in a decreasein the initial weight of the polymer film to 50% after15 days (25 days in total).

The water uptake behavior is shown in Fig. 4. Incontrast to the weight loss, mesh size had a majorinfluence on the uptake of water, for the same Fig. 3. Influence of mesh width and polymer properties on thepolymer film (50:50 PLGA, RG503). Generally, this degradation behavior: upper, RG503 films on 720 and 50 mmmatrix exhibits three distinct phases of degradation, meshes; lower, 720 mm meshes with R207, RG858 and RG503

polymer films. Mean values and standard deviations are shownbeginning with an initial uptake of ,3% by weight,(n54).followed by a plateau phase of no uptake (one to

seven days) and, finally, an uptake of water at a rateof 4–5% per day, which reached maximum water The decrease in the molecular weight of the matrixcontents of 20 to 25% at day 14. Subsequently, the is shown in Fig. 5. Starting from the same initialwater content decreased continuously for both mesh molecular weight, PDLLA (R207) degraded con-sizes. Optically, none of the specimens showed signs siderably slower than 85:15 PLGA (RG858), being

5 4of macroscopic deterioration, e.g. loss of membrane 1.2?10 and 6?10 g/mol after 25 days, respectively.segments. The 85:15 PLGA films revealed only a The 50:50 PLGA (RG503) matrix exhibited ex-slow rate of water uptake after an initial absorption ponential degradation characteristics, reaching aof 2% during the first 24 h. molecular weight of below 580 g/mol in 25 days.

Table 3Influence of the mesh size on the characteristics of the film, i.e., weight gain, film thickness and extension of the meniscus between the filmand the fiber

Mesh size Relative weight grain Film thickness Relative meniscus area(mm) (%) (mm) (%)

50 80 50 30200 50 50 10720 40 60 10

86 J. Mayer et al. / Journal of Controlled Release 64 (2000) 81 –90

Fig. 5. Time-dependent decrease in the molecular weight (Mw) of50:50 PLGA (RG503) films on 720 and 50 mm meshes (upper)Fig. 4. Relative water content of different combinations ofand of PDLLA (R207) and 85:15 PLGA (RG858) (lower). Meanpolymer matrices and woven fabrics. Upper, 50:50 PLGAvalues and standard deviations are shown (n53).(RG503) films on 720 and 50 mm meshes; lower, 720 mm meshes

with R207, RG858 and RG503 polymer films. Mean values andstandard deviations are shown (n54).

The presence of PVLA (at low and high concen-trations) seemed to slightly accelerate the rate of

Generally, the mesh size had no significant effect on aggregate formation. After 48 h, almost all cells onthe decrease in molecular weight. PVLA-coated surfaces had aggregated, while aggre-

The results of the preliminary hepatocyte culturing gate formation on blank PLGA surfaces was not yetstudy are shown in Fig. 6. Hepatocytes attached onto finished.the PVLA- and EGF-free PLGA surfaces without Neither the aggregation kinetics nor the aggregateforming aggregates. On PVLA-coated, EGF-free size depended on the initial distribution of cells inscaffolds, the hepatocytes showed a different mor- the presence of EGF. A high concentration of cellsphology without the initial PVLA concentration per mesh would result in the formation of two orhaving an influence. Generally, cells were more more identical spheroids. The mesh size had anspread out and very few single cells were observed. influence on the formation kinetics and on the shapeAfter two days, hepatocytes showed many short of the aggregates. In the case of 200 mm meshes, themembrane extensions (lamellipodia), which led to cells were often observed to spread throughout thedirect cell–cell coupling and resulted in the forma- pore. In fabrics with small pore sizes (50 mm), thetion of multicellular aggregates after the fourth day. hepatocytes initially covered the fibers as well. AfterIN the presence of EGF (50 ng/ml), the hepatocytes four days, large aggregates were observed that wereformed large multicellular aggregates within 48 h. anchored in several pores simultaneously. Hepat-

J. Mayer et al. / Journal of Controlled Release 64 (2000) 81 –90 87

Fig. 6. Influence of mesh size and coating on hepatocyte aggregation. (a) No PVLA coating, no EGF, 720 mm mesh size. After 96 h,hepatocytes had a round shape. Magnification, 1003. (b) PVLA coating (1 mg/ml) and no EGF, 720 mm mesh size. There was directcell–cell coupling, which resulted in the formation of multicellular aggregates on the fourth day. At 100 mg/ml PVLA, no differences wereobserved. Magnification, 1003. (c) No PVLA coating and 50 ng/ml EGF, 720mm mesh size. All hepatocytes within a mesh width wereinterconnected by large membrane extensions and formed a cellular network after 24 h. Magnification, 1003. (d) 1 mg/ml PVLA coating,50 ng/ml EGF, 720 mm mesh size. Spheroidal aggregates formed after 48 h. The PVLA coating had no effect. The diameter of theaggregate did not exceed approximately 160 mm. Magnification, 1003. (e) 100 mg/ml PVLA coating, 50 ng/ml EGF, 200 mm mesh size.Hepatocytes had large extensions of the cell membrane and were attached on different wires of the mesh simultaneously. Magnification,1003. (f): 100 mg/ml PVLA coating, 50 ng/ml EGF, 720 mm mesh size. SEM image showing aggregate formation.

88 J. Mayer et al. / Journal of Controlled Release 64 (2000) 81 –90

ocytes on 720 mm scaffolds were observed to form in vitro degradation varied considerably, dependingspheroidal aggregates within 48 h, whereas aggregate on polymer and fabric. The water uptake characteris-formation took more then 72 h on 200 mm and 50 tics of scaffolds with a 50:50 PLGA matrix (RG503)mm scaffolds. In large meshes, aggregates showed a is closely correlated to the weight loss data (Fig. 3).smooth spheroidal morphology, while they appeared For the first 14 days, the relative water content of themore irregular in smaller meshes. This can be films tends to be directly proportional to the rate ofexplained by the number of cells involved in the weight loss and is independent of the fabric’sformation of aggregates: in the large pores, more geometry. The behavior of the other polymers duringcells are available and the aggregates become larger the late phase of their degradation should not beand, thus, more sphere-like. compared to RG503 because of their slow degra-

dation.The experiments described, in which primary

4. Discussion mouse hepatocytes were seeded onto differentcomposite scaffolds, were performed as a prelimin-

The two-phase casting technique allows one to ary study. Hepatocytes in spheroidal aggregatescoat woven fabrics on one side with a biodegradable typically exist in a differentiated, nonproliferativepolymer film. This system exhibits a polarized state, whereas those cells that develop into mono-structure, with a plain ‘film side’ and a highly layers are more likely to proliferate [22]. Theorganized ‘fabric side’, with well-defined pores of kinetics of spheroid formation has been correlatedbetween 800 and 50 mm in width, depending on the with the interactions between the contractile strengthselected mesh size of the fabric. and the cell–substrate adhesion strength of hepat-

Weight loss of biodegradable, bulk-eroding poly- ocytes [22]. Therefore, strong cell–substrate adhe-mers begins when a critical molecular weight has sion should result in monolayered cultures.been reached and the degradation products have We used blank and PVLA-coated PLGA/PETbecome soluble. Our experiments generally demon- scaffolds to investigate the differences in non-spe-strate an initial weight loss of 2% after two days. cific (on the PLGA) and receptor-mediated (byThis might be related to diffusion of residual solvent asialoglycoprotein receptors on the PVLA) cell at-and of soluble oligomers that originated from the tachment on our composite scaffolds. Spheroidalfilm preparation. The degradation characteristics did aggregate formation was observed in all cases,not depend on the supporting fabric, as one might however, in the absence of PVLA, the spheroidspeculate from the relationship between mesh size formation was slightly retarded. In the case ofand specific surface area. PLGA, the hepatocytes may have attached using

The pronounced weight loss of the 50:50 PLGA hydrophobic and ionic interactions of the cell mem-matrix after eight days can be related to the molecu- brane and the substrate. Based upon our observa-lar weigh,t which was reduced to less than 8000 tions, we assume that the adhesion strength ofg /mol and, thus, indicates that the polymer is hepatocytes to the PLGA substrate is, although non-partially soluble. The increased scatter in molecular specific, within the same range as for PVLA. Inweight data at a mesh size of 50 mm can be contrast to Ref. [22], no concentration effects couldexplained by the larger contribution of the meniscus be detected, however, this should be interpreted witharea to the film (see Table 3). Furthermore, due to care since the water absorption of the polymer doesthe smaller open area of the 50 mm fabric, a larger not allow quantitative determination of PVLA ad-proportion of the matrix covers the threads. This sorption.might indicate that the degradation behavior of the Generally, it was observed that, independent of thefilm is changed where it forms an interface with the cell density, all hepatocytes within a mesh werethreads of the fabric. However, the differences in interconnected by large membrane extensions (in theopen area among the fabrics are too small to allow presence of EGF) and formed a cellular networkquantitative conclusions to be drawn. after 24 h. In scaffolds with large fabrics (720 mm),

The relative water content of the scaffolds during spheroidal aggregates of |160 mm formed within 48

J. Mayer et al. / Journal of Controlled Release 64 (2000) 81 –90 89

h. We assume that this efficient aggregate formation support the assumptions that our scaffold was basedis caused by the scaffold topography. The fibers of upon.the large-sized fabric (wire diameter, 350 mm) divide In this study, we showed the importance of matrixthe surface into well defined compartments, from engineering in achieving specific cell functionality inwhich, the cells cannot migrate. Therefore, the tissue engineering, e.g. hepatocyte aggregation beingnumber of cells that can interact is determined by the influenced by the three-dimensional structure of theconcentration of cells and the scaffold geometry. In scaffold. According to the criteria of structuralsmaller meshes, we observed cell–cell interactions biocompatibility, spatial reorganization of livingacross the fibers, which slowed down the formation tissue by appropriate three-dimensional scaffold de-of aggregates. In the case of 50 mm scaffolds, we sign is a precondition to achieving highly differen-could observe aggregation of hepatocytes covering tiated and functional cell cultures. Tissue engineeringseveral mesh areas. Therefore, even from these is considered as one of the key technologies in thepreliminary data, it can be concluded that the geome- development of new biomaterials and of new thera-try of the scaffold plays a significant role by peutic opportunities.imposing physical barriers to cell migration andcell–cell interaction. Furthermore, co-culturing ofhepatocytes is necessary for the development of Acknowledgementsliver-like tissues in vitro [22,24,28]. The proposedscaffold system is expected to offer the possibility of

The authors would like to express their gratitudeco-culturing hepatocytes and non-parenchymal cells;

to H. Ise, S.-H. Kim and R. Takei at the Faculty ofhepatocytes on one side and non-parenchymal cells

Bioscience and Biotechnology, TIT Tokyo and to B.on the other side of the membrane. Using a porous

Eder-Aebersold, S. Ritter, K. Ruffieux and R. Reberdegrading membrane, it may be possible to create an

at the Chair of Biocompatible Materials Science and‘‘artificial Disse’s space’’, where ECM proteins that

Engineering, ETH Zurich. Financial support fromare excreted by the non-parenchymal cells slowly

both the Tokyo Institute of Technology and thereplace the degrading polymer membrane.

Swiss Federal Institute of Technology is greatlyacknowledged.

5. ConclusionReferences

This study concentrated on stable PET fabrics that[1] E. Wintermantel, S.-W. Ha, Biokompatible Werkstoffe undwere coated with various PLA/PGA polymer ma-

¨Bauweisen. Implantate fur Medizin und Umwelt, Springer,trices, in order to cover a wide range of degradationBerlin, 1996, p. 6.rates. The investigation of the in vitro degradation of

¨[2] E. Wintermantel, J. Mayer, J. Blum, K.-L. Eckert, P. Luscher,our scaffolds has shown that degradation begins at M. Mathey, Tissue engineering scaffolds using superstruc-the film/fabric interface. The system maintained it’s tures, Biomaterials 17 (1996) 83–91.structural integrity and no detachment of the film [3] E. Wintermantel, A. Bruinink, K.-H. Eckert, K. Ruffieux, M.

Petitmermet, J. Mayer, Tissue engineering supported withoccurred.structured biocompatible materials: goals and achievements,Primary mouse hepatocytes were cultured on thein: M.O. Speidel, P. Uggowitzer (Eds.), Materials in Medi-

scaffolds in different media and under different ECM cine, 1998, pp. 1–131.conditions. Hepatocytes attached on the PLGA sur- [4] S. Kasai, M. Sawa, M. Mito, Is the biological artificial liverfaces, but no aggregation was observed in the clinically applicable? A historic review of biological artificial

liver support systems, Artif. Organs 18 (1994) 348–354.absence of EGF. The influence of PVLA coatings on[5] V. Dixit, Development of a bioartificial liver using isolatedthe aggregation behavior could not differentiated

hepatocytes, Artif. Organs 18 (1994) 371–384.decisively. The different mesh sizes had an influence [6] P.S. Malchesky, Nonbiological liver support: historic over-on the kinetics of aggregate formation. These results view, Artif. Organs 18 (1994) 342–347.have to be judged as preliminary, however, they [7] T.J. Yu, C.C. Chu, Biocomponent vascular grafts consisting

90 J. Mayer et al. / Journal of Controlled Release 64 (2000) 81 –90

of synthetic absorbable fibers. I. in vitro study, J. Biomed. [19] A.G. Mikos, Y. Bao, L.G. Cima, D.E. Ingber, J.P. Vacanti, R.Mater. Res. 27 (10) (1993) 1329–1339. Langer, Preparation of poly(glycolic acid) bonded fiber

[8] A. Kamlot, J. Rozga, D. Watanabe Frederick, A. Demetriou structures for cell attachment and transplantation, J. Biomed.Achilles, Review: artificial liver support systems, Biotechnol. Mater. Res. 27 (1993) 183–189.Bioeng. 50 (1996) 382–391. [20] D.J. Mooney, D.F. Baldwin, N.P. Suh, J.P. Vacanti, R.

[9] V. Dixit, E. Pishkin, M. Arthur, A. Denizli, S. Tuncel, E. Langer, Novel approach to fabricate porous sponges poly-Denkbas, G. Gitnick, Hepatocyte immobilisation on PHEMA (DL,-lactic–co-glycolic acid) without the use of of organicmicrocarriers and its biologically modified forms, Cell solvents, Biomaterials 17 (1996) 1417–1422.Transplant. 1 (1992) 391–399. [21] D.J. Mooney, K. Sano, P.M. Kaufmann, K. Majahod, B.

[10] D.J. Mooney, S. Park, P.M. Kaufmann, K. Sano, K. McNa- Schloo, J.P. Vacanti, R. Langer, Long-term engraftment ofmara, J.P. Vacanti, R. Langer, Biodegradable sponges for hepatocytes transplanted on biodegradable polymer sponges,hepatocyte transplantation, J. Biomed. Mater. Res. 29 (1995) J. Biomed. Mater. Res. 37 (1997) 413–420.959–965. [22] A. Kobayashi, M. Goto, T. Sekine, A. Masumoto, N.

[11] E. Wintermantel, L. Cima, B. Schloo, R. Langer, Angiopo- Yamamoto, K. Kobayashi, T. Akaike, Regulation of differen-larity: a new design parameter for cell transplantation and its tiation and proliferation of rat hepatocytes by lactose-carry-application to degradable systems, Trans. Am. Soc. Artif. ing polystyrene, Artif. Organs 16 (1992) 564–567.Intern. Organs, July /September, 1991. [23] T. Akaike, M. Goto, H. Yura, A. Kobayashi, C. Cho, A.

[12] A.D. Moscioni, G. Backfisch, T. Bellew, D.L. Black, A.A. Maruyama, K. Kobayashi, New frontier of biomimeticDemetriou, Long-term cryopreserved human hepatocytes glycoengineering for cellular and tissue engineering, in: N.maintain drug metabolizing ability, Surg. Forum 41 (1990) Ogata, S.W. Kim, J. Feijen, T. Okano (Eds.), Advanced3–4. Biomaterials in Biomedical Engineering and Drug Delivery

[13] A. Cai, Z. Shi, M. Sherman, A.M. Sun, Development and systems, Springer, Tokyo, 1996, pp. 173–178.evaluation of a system of microencapsulation of primary rat [24] B.A. Naughton, J.S. Roman, B. Sibanda, J.P. Weintraub, V.hepatocytes, Hepatology 10 (1989) 855–860. Kamali, Stereotypic culture systems for liver and bone

[14] C.F.W. Wolf, B.E. Munkelt, Bilirubin conjugation by an marrow: evidence for the development of functional tissue inartificial liver composed of cultured cells and synthetic vitro and following implantation in vivo, Biotechnol. Bioeng.capillaries, Trans. Am. Soc. Artif. Intern. Organs 21 (1975) 43 (1994) 810–825.16–26. [25] K. Taguchi, M. Mastushita, M. Takahashi, J. Uchino,

[15] F.J. Wu, M.V. Peshwa, F.B. Cerra, W.S. Hu, Entrapment of Development of a bioartificial liver sandwich — culturedhepatocyte spheroids in a hollow fiber bioreactor as a hepatocytes between two collagen gel layers, Artif. Organspotential bioartificial liver, Tissue Engineering 1 (1995) 20 (1996) 178–185.29–40. [26] R. Takei, K.H. Park, H. Ise, A. Maruyama, Y. Ebara, Y.

[16] D.J. Mooney, C.L. Mazzoni, C. Breuer, K. McNamara, D. Okahata, T. Akaike, Coated antireceptor antibody as anHern, J.P. Vacanti, R. Langer, Stabilized polyglycolic acid extracellular matrix for liver tissue engineering, Tissuefibre-based tubes for tissue engineering, Biomaterials 17 Engineering 3 (1997) 281–288.(1996) 115–124. [27] P.O. Seglen, Preparation of isolated rat liver cells, Methods

[17] L. Grizzi, H. Garreau, M. Vert, Hydrolytic degradation of Cell Biol. 13 (1976) 29–83.devices based on poly(DL-lactic acid) size dependence, [28] T. Akaike, S. Tobe, A. Kobayashi, M. Goto, K. Kobayashi,Biomaterials 16 (1996) 305–311. Design of hepatocyte-specific extracellular matrices for

[18] C.M. Agrawal, G. Niederhauer, K.A. Athanasiou, Fabrica- hybrid artificial liver, Gastroenterol. Jpn. 28 (1993) 45–52.tion and characterization of PLA–PGA orthopedic implants,Tissue Engineering 1 (1995) 241–252.