Cell seeding in porous transplantation devices

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Cell seeding in porous transplantation devices Heidi L, Waldf, Georgios Sarakinos*, Michelle D. Lyman, Antonios G, Mikoss, Joseph P. Vacanti** and Robert Langer Department of Chemical Engineering, *Department of Chemistry, Massachusetts Institute of Technology, Massachusetts Avenue, Cambridge, MA 02139, USA; **Department of Surgery, The Children’s Hospital, Harvard Medical School, 300 Longwood Avenue, Boston, MA 02115, USA

77

Porous laminated discs of 1.35 cm diameter and thickness of 0.5 cm fashioned from biodegradable polymers were used as scaffolds for the transplantation of isolated cell populations. The distribution of cells seeded in these devices via injection was modelled with a system of dyed polymeric microparticles. Optimization of parameters related to device design and surgical injection conditions was carried out to maximize the device volume effectively employed in cell transplantation. The area of distribution on the top surface of each device was determined by image analysis techniques and used as a measure of the spatial distribution of injected particles. For poly(L-lactic acid) devices of porosity of 0.83 and median pore diameter of 166pm seeded with 6pm beads under standard injection conditions, the average surface area of distribution was 44.45% (f3.36%). The device pore size was found to be a crucial determinant of particle distribution, whilst particle size in the range of l-10pm was not found to be important for the devices tested. Application of these results to the seeding of hepatocyte suspensions was made. Keywords:

Cell transplantation,

celi seeding,

biodegradable

polymer,

poly(L-lactic

acid)

device

Received 1 May 1992; accepted 17 July 1992

Cell transplantation has been explored as an alternative to various means of replacing tissue function’-4. Using this approach, individual cells are harvested from a healthy section of donor tissue, isolated and implanted in the patient at a desired site. Cell transplantation has several advantages over whole organ transplantation. Because the isolated cell population can be expanded in vitro using cell culture techniques, only a very small number of donor cells are needed to prepare an implant. Consequently, the living donor need not sacrifice an entire organ. The use of isolated cells also allows removal of other cell types which may be the target of immune responses, thus diminishing the rejection pmcess5. In addition, major surgery on the recipient and donor and its inherent risks are avoided. Finally, the cost of the procedure may be significantly reduced. Isolated cells cannot form new tissues on their own. They require specific environments which very often include the presence of supporting material to act as a scaffolds are template for growth 6,7. Three-dimensional Correspondence to Professor R. Langer. +Present address: Harvard Medical School, 25 Shattuck Street, Boston, MA 02115, USA. BPresent address: Department of Chemical Engineering and The Institute of Biosciences and Bioengineering, Cox Laboratory for Biomedical Engineering, Rice University, PO Box 1892, Houston, TX 77251, USA.

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used to mimic their natural counterparts, the extracellular matrices (ECM) of the body. They serve as both a physical support and an adhesive substrate for isolated parenchymal cells during in vitro culture and subsequent implantation3’ 4. Because of the multiple functions these materials must fulfil, the physical and chemical requirements are numerous. To accommodate the adhesion of the largest number of cells, a cell transplantation device must have a large surface area for cell adhesion. High porosity provides adequate space for cell seeding, growth and ECM production. A uniformly distributed and interconnected pore structure is important so cells are easily distributed throughout the device, and an organized network of tissue constituents can be formed. In the reconstruction of structural tissues like bone and cartilage, tissue shape is integral to function, Therefore, these scaffolds must be processable into devices of varying thickness and shape. Furthermore, due to eventual implantation, the scaffold must be made out of a biocompatible material. As the transplanted cell population grows and the cells function normally, they will begin to secrete their own ECM support. The need for an artificial support will gradually diminish, and so a biodegradable implant will be eliminated as its function is replaced. Efforts in our laboratory have led to the development 0

1993 Butterworth-Heinemann Ltd 0142-9612/93/040270-09

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of a process to make transplantation devices from biodegradable polymers which satisfy the above requirements’. The potential for growth and normal organization of hepatocytes on such supports to regenerate liver function has been demonstrated” lo. Clinical implementation of this method remains limited by the underutilization of the device pore volume and incomplete knowledge of the optimal conditions for surgical administration of the cell population into the devices. Here, a series of studies is presented to elucidate the methods by which to achieve a uniform distribution of cells throughout transplantation devices. In this manner we can maximize the utilized volume of the device, take full advantage of the shape of the device, and ensure that the entire tissue is uniformly cell-rich. If all cells grow and produce ECM macromolecules at the same rate, then a uniform section of functioning tissue has been created.

MATERIALS AND METHODS Polymer

foams

Polymer foams were prepared by a solvent-casting technique developed by Mikos et al.‘. Briefly, 0.5 g of solid poly(L-lactic acid] (PLLA) (Polysciences, Warrington, PA, USA) were dissolved in 8 ml of chloroform at room temperature. Sodium chloride crystals (4.5 g, Mallinckrodt, Paris, KY, USA] which were sieved to the desired particle size range (106-150 or 250-500pm) were added to the polymer-solvent solution. The mixture was vortexed until a uniform suspension was obtained and cast on to a glass petri dish 5 cm in diameter. The chloroform was evaporated from the covered dish at room temperature for 48 h. Residual solvent was removed by placing the membrane under vacuum (13.3 Pa) for 24 h. Removal of salt particles was accomplished by leaching in water for 36 h. During this time period, the water was changed every 6 h. The polymer foams were dried and stored desiccated until use. The pore size range of the resultant polymer membranes showed a direct relationship to, but was not exactly equal to, the salt particle size range used”. The median pore diameters of the produced membranes were determined by mercury porosimetry’ to be 126pm using salt particles of size in the range of 106-150pm and as 166pm with those of 250-500pm. The membrane porosities were also measured as 0.80 and 0.83 for the small and the large particles, respectively.

Fabrication

of cell transplantation

devices

Three-dimensional devices were constructed in a range of configurations by a membrane lamination method’. A catheter was inserted into the centre of each device as a route for injection of cells into the bulk of the polymer. The steps involved in preparation of the devices are illustrated in Figure 1. Discs of 1.35 cm in diameter were cut from the polymer foams using a cork borer. The discs were glued to one another using small amounts of chloroform to wet adjacent surfaces. The devices consisted of three layers of PLLA foams. A 5 cm piece of medical grade silicone tubing (0.03 inches inner diameter and 0.065 inches outer diameter: American Scientific Products, McGaw Park, IL, USA) which served as the catheter was incorporated into the device across the

Figure 1 Schematic diagram of the procedure for assembling cell transplantation devices from three polymer foams and a catheter with two holes cut out of it. The device is shown from the top view and in cross-section at each stage of assembly.

diameter of the middle layer. A knot was tied at one end of the tubing. At a distance of 0.675 cm from the knot, two identical rectangular holes of length l/l6 inches were opened with a razor blade for bead injection. (The desired hole size was cut out with the aid of an English hexagonal key of size 0.035 inches which was put _ Biomaterials

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through the tubing.) The tubing was positioned in such a way that the two holes were at the centre of the device and faced the edge. The completed device was lyophilized for 24 h to remove residual chloroform. The devices to be used in cell experiments were sterilized using ethylene oxide gas for 12 h. A photograph of a completed device is shown in Figure 2.

Hepatoc~e

harvest

Hepatocytes were harvested from Fisher rats (Charles River Breeding Laboratories, Wilmington, MA, USA) according to the technique of Seglen as modified by Cima et aL1’. Each rat was anaesthetized using methoxyflurane (Pitman-Moore, Washington Crossing, NJ, USA) and its abdomen was shaved, prepared with betadiene and opened under sterile conditions. The liver was isolated, heparinized with 100 units heparin (ElkinsSinn, Cherry Hill, NJ, USA), and the portal vein was cannulated with a Z&gauge plastic i.v. cannula (Critikon, Tampa, FL, USA). The inferior vena cava was transected, and the liver was flushed with 2 ml of sterile saline. The liver was then transferred to a sterile dish and perfused with an oxygenated solution of 0.025% collagenase, Type II (Appel Products, West Chester, PA, USA). The perfusion occurred for 20 min after which cells were dispersed in culture medium, William’s E with 10 ngiml EGF (Collaborative Research, Bedford, MA, USA), 20 munitsfml insulin (Gibco, Grand Island, NY, USA], 5 nM dexamethasone [Sigma, St Louis, MO, USA), 20 mM pyruvate (Gibco) and 100 units/ml penicillinstreptomycin (Gibco). The debris was removed by centrifugation and washing in the culture medium. The cell viability at plating was 86% as determined by direct cell counts using trypan blue (Sigma).

Microbead

injection

experiments

Dyed [red) monodisperse microparticles of polystyrene (Polysciences) ranging in size from 1 to 10 ,um in diameter were injected into the transplantation devices. The devices were thoroughly wet before injection in ethanol for 10 min and then in 0.9 wt% saline solution for 10 min.

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The microparticle suspension was injected from a syringe through a 12 inch catheter and al-gauge needle (Minicath Infusion Set, Deseret, Sandy, UT, USA) and into the device’s own catheter using a Harvard pump. Figure3 shows a schematic of the apparatus. The device was then sectioned and photographed with a zoom macroscope (model M420, Wild Heerbrugg, Heerbrugg, Switzerland). Most of the devices were photographed when still wet. The surface area measurements of devices which were photographed after drying for 24 h were compared only to other dry devices. The basic injection conditions used for testing were an injection rate of 0.5 ml/min, 6 pm in diameter beads and catheter holes (l/l6 X l/16 inch’) at the centre of devices with a median pore diameter of 166 pm. Total injection volume was 0.5 ml. Variables included the rate of injection of the microbead suspension, the bead diameter, the catheter hole position and size, the device pore size and the device thickness. Triplicates were run for each variable tested. Acid orange 8 dye (0.5%) (Sigma) was injected into some of the devices in a similar manner.

Hepatocyte! injection

and histology

A suspension of 1.0 X lo7 cells/ml was injected into each device in a procedure identical to that used for microbead injection. The total injection volume was 0.5 ml. The cellular experiments were pe~ormed under sterile conditions. Samples were prepared for histology by removing the culture medium from the sample dish and fixing the device in 3% glutaraldehyde (Polysciences) for 15 min. The samples were rinsed and stored in 10% neutral buffered formalin solution (Sigma) until sectioning and staining. The samples were sliced into thin sections and stained with haematoxylin and eosin (H&E) which allowed for visualization of cells and cell nuclei.

Image analysis Image analysis was performed using a Magiscan 2 image analysis system (Joyce-Loebl, Tyne and Wear, UK) equipped with a Polaroid MP-4 Land camera. Polaroid

19

I.1

! 3

Figure 2 Photograph of a complete transplantation device. The catheter is partially removed from the device so that its holes are visible. I-- ._.._..--. _~~~~~~-~--.--..-._ _ _~_._~--_.--.Biomaterials

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- .. . .-.___.

Cell seeding - in porous

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Figure 3 Schematic diagram of the experimental for microbead injection into porous devices.

H. L. Wald et al.

apparatus

photographs of the devices at 7.875 magnification were analysed for the percentage of the top surface area of 1.43 cm2 which was coloured red by the presence of dyed microspheres. Each photograph was analysed by interfacing the Polaroid camera to the black and white video monitor of the image analysis system. The greyscale of the image was segmented [Figure 4) on the video screen. Any part of the image that was dark (i.e. red from the dyed microspheres) was coloured white. The detected area of each white region was determined by the number of non-zero pixels which constituted the feature region. The sum of the detected areas was recorded as the total area of dyed microsphere distribution for the top surface of the device.

RESULTS Microbead distribution throughout the device volume was measured by the area of bead distribution on the external surface of the device. It was this surface which was furthest away from the injection site; it therefore represented the distribution of microbeads at the most distant site from where the microbeads were injected. This surface area was also used because it was possible to photograph it without disrupting the device. Figure 5 contains photographs of a representative sample of devices seeded with the dyed microspheres. Measurable differences in surface area of coverage were found for devices which were photographed while still wet versus those photographed while still dry (Figures 5b,d).

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Figure4 Schematic diagram of a device after injection of dyed microbeads. Before segmentation of grey scale (top) and after segmentation (bottom). The surface coverage was determined as the percentage of the whiti? region.

Variations in the microbead distribution were observed for many of the factors examined. Determination of statistical significance was made using the Student’s t test for unpaired samples. These results are summarized below.

Prewetting

devices

Injection of liquid or microsphere suspension into dry devices was not useful for the purpose of seeding. Figure 6 shows two devices, both injected with acid orange 8 dye. The dry device prohibited entry of the dye solution into its porous structure. The device which was prewetted with ethanol allowed the rapid transport of dye into the pores. All further evaluation was completed on devices which were prewetted according to this protocol.

Device uniformity Devices injected either with acid orange 8 dye or with microparticle suspension were studied using device cross-sections and subjected to visual examination. It was observed that the pores of adjacent layers of the laminated devices were well interconnected, evidenced by the rapid and unhindered transport of both dye solution and model beads across the interface of adjacent layers. Thicker devices were constructed by adding an additional porous layer to the top and bottom surfaces of Biomalerials

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Figure 5 Photomicrographs of the top view of PLLA devices fabricated with salt particles of size in the range of 106-150pm, a, and 250400pm, b-d. The devices were seeded by a suspension of model microbeads of 6pm diameter. Both catheter holes were placed at the centre of the devices, a,b,d; for c, one catheter hole was positioned at either end of the device. The photomicrographs, a,b, were taken with wet devices and c,d with dry devices (scale bars = 0.5 cm).

Figure 6 Photomicrographs of cross-sections of PLLA devices injected by acid orange 6 dye solution. Device a was injected dry whereas device b had been prewetted in ethanol (scale bars = 0.5 cm). Biomaterials

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Figure 7 Photomicrograph of the cross-section of a PLLA device with five layers seeded by a suspension of model microbeads of 6pm diameter (scale bar = 0.5 cm).

each device. Figure 7 is a cross-section through one of these devices. Materials Other materials of the poly(DL-lactic-co-glycolic acid) family (PLGA) were tested to ascertain the usefulness of membrane lamination on devices fashioned from these materials. Two copolymers of 85/15 and 50/50 molar ratios of lactic acid to glycolic acid were used (Medisorb, Cincinnati, OH, USA). Mic~sphe~s injected under standard conditions into devices made from PLLA showed an average surface area of distribution of 44.45% (f3.36%), whilst those made from PLGA 85/15 and PLGA SO/50 were covered 59.52% (+4.20%) and 51.09% (+4.08%), respectively (Figure 8). The variation of the 100

Figure 13 Photomicrograph from the histological section of a PLLA device seeded by a hepatocyte suspension. Several degenerated hepatocyte forms are observed. surface coverage with the polymer type was explained by the different thickness of the devices*made of these polymers. The thickness of devices prepared using the identical technique was 4999 (+72)pm for PLLA, 3531 (+427)pmforPLGA85/15 and4484(f298)hmforPLGA 50/50. The values of device thickness are averages ks.d. of five measurements. As the device thickness increased, the number of particles reaching the outer surface diminished resulting in smaller surface coverage. All further evaluation was performed on devices of PLLA. Catheter

80

T T

20

-I

0 PLLA

PLCA 85115 Polymer

Figure8 Surface coverage of devices of different polymers seeded by a suspension of model microbeads of diameter 6pm at a rate of 0.5 mllmin. The reported values correspond to averages rts.d. of three experiments. The devices were examined wet.

hole size and placement

Two different configurations of catheter holes were tested and are shown in Figums 5c,d. The variation in the surface coverage is presented in Figure 9, The first configuration had two holes placed 1 cm apart and on opposite sides of the catheter. The second configuration had both holes at the centre of the device. Microparticles injected through the catheter of the first configuration left the catheter through the hole closest to the catheter entrance to the device. They did not reach the second hole under experimental injection conditions (0.5 ml/m@. In addition, there was much particle leakage out of the edge of the device nearest to the first hole. In this case, particle distribution was limited to an approximately spherical volume near the first hole covering only 18.75% (+X04%) of the surface (photographed dry). Devices of the second configuration showed a much larger distribution of particles radially outward from the centre of the device, and leakage from the edges was drastically reduced. The surface area of coverage was 28.71% (+2.99%) of the total (dry). The difference in mean surface area of coverage using the two catheter configurations was statistically significant (P = 0.012). The second Biomaterials

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_-I20

El-El 20

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0 1116” Centre wet

2116” Centre wet Hole size

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I

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:/,‘d”:

dry

dry

position

Figure 9 Variation of the surface coverage of PLLA devices seeded by a suspension of model microbeads of diameter 6 pm at a rate of 0.5 ml/min with the size and the position of the catheter holes. The reported values correspond to averages +s.d. of three experiments. The devices were examined as indicated.

I

0.25

0.50 Injection

speed,

1 .oo mllmin

Figure 10 Variation of the surface coverage of PLLA devices seeded by a suspension of model microbeads of diameter 6 ,um with the suspension injection speed. The reported values correspond to averages fs.d. of three experiments. The devices were examined wet.

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configuration was used in the remainder of the devices tested. Devices with catheter holes 2116 inches in length showed a uniform bead distribution over 54.87% (+8.58%) of the surface. Devices with the standard hole size of l/16 inches also had a uniform bead distribution covering an average of 44.45% (+3.36%) of the surface area (Figure 9). The reported difference in the mean values is not significant (P = 0.215).

Injection

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r

tl_l 1 T

rate

Trials utilizing the fastest injection rate (1.0 ml/min) showed a mean detected surface area of 46.45% (f4.53%) of the whole, and even dispersion over most of the surface area (Figure 10). The slowest injection rate (0.25 ml/min) showed poorer average distribution of 35.71% (f8.38%) with the beads remaining in the centre of the device. Although these data showed an increasing surface area of distribution as injection rate increases, the mean values were not statistically different from the standard case (0.5 mbmin).

-r

20

0

-

166

126

Median

pore

diameter,

urn

Figure 11 Variation of the surface coverage of PLLA devices seeded by a suspension of model microbeads of diameter 6pm at a rate of 0.5 ml/min with the foam median pore diameter. The reported values correspond to averages fs.d. of three experiments. The devices were examined wet.

Particle size and pore size Microspheres injected into devices with a median pore diameter of 166 pm were dispersed in a spherical volume about the centre of the device where the catheter holes were located (Figure 21). The variation in the surface coverage is shown in Figure 11. For these devices, as previously mentioned, the percentage surface area of distribution was 44.45% (f3.36W). In the devices with median pore diameter of 126,~m, the surface area of distribution of beads was diminished to 28.76% (*6.18%). This difference in distribution is statistically significant Biomaterials

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(P = 0.007). Very little radial dispersion of the beads was observed in the 126.um devices. Instead, an axial channeling along the length of the catheter and a high degree of microparticle leakage from the device at both the entrance and exit of the catheter were observed. Monodisperse beads of 1 and 10pm in diameter did not show any difference in the average surface area of distribution (Figure 22). The 10 pm beads, were, however, distributed at the device surface in a patchy manner.

Cell seeding

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devices:

_

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6 Microbead

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diameter,

urn

Figure 12 Variation of the surface coverage of PLLA devices seeded by a suspension of model microbeads at a rate of 0.5 ml/min with the microbead diameter. The reported values correspond to averages +s.d. of three experiments. The devices were examined wet.

Ex viva experiments Hepatocytes injected into PLLA transplantation devices [with median pore diameter of 166pmm) under standard conditions (0.5 ml/min) were well distributed throughout the devices as indicated by the colouring of the external surfaces. Thin sections revealed numerous hepatocytes at cross-sections taken at distances of 114 inches from the centre of the devices both parallel and perpendicular to the catheter. Some of the hepatocytes appeared to be injured (Figure 13). All were spherical in shape, and none were flattened out along the implant material.

DISCUSSION The model microparticles proved to be a fast, simple way in which to determine the parameters that are important for seeding cells in porous devices. Because isolated cells are spherical in suspension, the microbeads are expected to provide a general idea of the behaviour of a cell suspension to a first approximation. This system allows for the optimization of injection conditions without the complication and waste of using a large and steady supply of animal cells. The use of dyed microspheres permitted rapid determination of particle distribution within the device. Polymeric microparticles were also used to examine the internal pore structure of the laminated devices. It was evident from cross-sections that the lamination process did not cause obstruction of the flow of fluids or microspheres between pores of adjacent layers (Figure 7). In thick devices constructed of five polymer membrane layers, the outermost layers were not adjacent to the catheter and showed dye and microsphere distribution consistent with the distribution in the rest of the device. The high hydrophobicity of PLLA prevented the entry of liquid into the air-filled porous structure of the dry

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PLLA devices. The prewetting procedure using ethanol was adopted to overcome this problem. Ethanol readily entered the porous structure of the device, after which it was readily displaced by water”. Factors which were found to be important in achieving uniform microbead distribution included the device pore size and the placement of catheter holes. The catheter hole placement had a marked influence on the bead distribution. When the holes were positioned near either end of the device, the beads did not reach the more distant half of the device. A great deal of bead leakage occurred using this configuration. Positioning both holes at the centre gave a much wider bead distribution. The results of our experiments also indicate that the pore size is an important determinant of microbead distribution because of the physical obstruction of flow which occursThe particle size has no effect on the distributionof particles with diameters in the range of l16pm injected into porous devices with median pore diameter of 166 pm. It is expected that for larger particles and/or devices with much smaller pores, the effect of the particle diameter becomes very significant. Nevertheless, we limited our studies to the above size range because it is typical of many mammalian cells. Because mammalian cells are fragile and easily disrupted by shear stresses, we tended to use injection rates that are as low as possible. It seemed likely, however, that higher flow rates would lead to better mixing and improved particle transport. The results of the microbead studies showed that as injection rate was increased over a small range, the microbead distribution in the devices increased in magnitude, but not by statistically significant amounts. The presence of damaged hepatocytes in the histological sections from the ex vivo experiments suggests that slower injection rates are preferable at the expense of marginal decreases in bead distributions. The use of polymeric microspheres allowed verification of the uniformity of the porous devices constructed by membrane lamination, It also proved useful in evaluating a number of parameters in device design. For the specific application of cell transplantation, this system permitted study of the effectiveness 01 various cell seeding techniques. The tortuous path presented to the cell by a porous scaffold restricts particulate transport and distribution within it. These experiments have demonstrated some of the factors which are critical for distributing the small particles throughout the devices. Some extrapolation to cell systems can be made. Efficient cell seeding is a crucial step in the use of threedimensional porous supports in cell transplantation. Elucidation of the factors which affect seeding will lead to a physiologically useful transplanted tissue.

ACKNOWLEDGMENTS Many thanks to Dr Magalie Fontaine for performing the hepatocyte perfusion and isolation, and to Dr Betsy Schloo of the Deborah Heart and Lung Center for the histology studies. This work was supported by a grant from Neomorphics, Inc. H.L. Wald would like to acknowledge the support of a National Science Foundation Graduate Fellowship. Biomaterials

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template

REFERBNCES Russell, P.S., Selective transplantation, Ann. Sum. 1985, 201,255-202 Yannas, I.V., Regeneration of skin and nerve by use of collagen templates, in Collagen ZZZ(Ed. M.E. Nimni), CRC Press, Boca Raton, FL, USA, 19813, pp. 87-116 Green, W.T.. Jr., Articular cartilage repair: behavior of rabbit chondrocytes during tissue culture and subsequent allo~aft~g, C&r. Orthop. 1977,124,237-250 Vacanti, J.P., Morse, M.A., Saltzman, W.M., Domb, A.J., Perez-Atayde, A. and Langer, R., Selective cell transplantation using bioabsorbable artificial polymers as matrices, Z. Pediatr. Surg. 1988, 23, 3-9 Schreiner, G.F., Flye, W., Brunt, E., Korber, K. and Lefkowith, J.B., Essential fatty acid depletion of renal allografts and prevention of rejection, Science 1988,240, 1032-1033 Vacanti, C.A., Lange& R., Schloo, B. and Vacanti, J.P., Synthetic polymers seeded with chondrocytes provide a

for new cartilage

devices: ML. Wald et a/. formation,

PZast. Reconstr.

Surg. 1991, 88, 753-759 7

8

9

10

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Freed, L.E., Marquis, J.C., Nohria, A., Emmanual, J., Mikos, A.G. and Langer, R., Neocartilage formation in vitro and in viva using cells cultured on synthe+ic biodegradable polymers,]. Biomed. Mater. Res. (in press) Mikos, A.G., Sarakinos, G., Leite, SM., Vacanti, J.P. and Langer, R., Laminated three-dimensional biodegradable foams for use in tissue engineering, BiomateriaJs (in press] Cima, L.G., Vacanti, J.P., Vacanti, C., Zngber, D.E,, Mooney, D. and Langer, R., Tissue engineering by cell transplantation using degradable polymer substrates, 1. Biomech. Eng. 1991,113,143-151 Cima, L.G., Ingber, D.E., Vacanti, J.P. and Langer, R., Hepatocyte culture on biodegradable polymeric substrates, Biotec~no~. Bioeng. 1991, 28,145-158 Miller, CA., Stability of interfaces, in Surface and Colloid Science (Ed. E. Matijevic), Vol. 10, Plenum Press, New York, NY, USA, 1978, pp 227-277

Biomedical Materials & Technologies

Bio~ompatibi~ty of Medical Devices Centro de Citologia Experimental, Porto, Portugal 28-30 June 1993 This course aims to give a comprehensive survey of current knowledge in the fields of biomateriais and biocompatibility, and will appeal to a multidisciplin~y audience. The st~c~re and properties of biomaterials will be reviewed, and particular attention paid to their uses in orthopaedics, dentistry and cardiovascular surgery. The biocompatibility of different materials will be fully examined. The interactions of host cells with biomaterials will be described and discussed, with emphasis being ptaced on the importance of biocompatibility to implant survival. For further information and registration details please contact: COMET’I’ Course Secretary, Department of Clinical Engineering, University of Liverpool, PO Box 147, Liverpool L69 3BX, UK. Fax: +44 051 706 5803

Biomaterials

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