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Stabilized polyglycolic acid fibre-based tubes for tissue engineering
Eiomateriak 17 (1996) 115-124 0 1996 Elsevier Science Limited Printed in Great Britain. All rights reserved 0142-9612/96/$15.00
ELSEVIER
Stabilized polyglycolic acid fibrebased tubes for tissue engineering D.J. Mooney *+$, C.L. Mazzoni*, C. Breued, K. McNamara*, D. Hem*, J.P. Vacanti’ and R. Langer* ‘Department of Chemical Engineering, Massachusetts institute of Technology, Cambridge, MA 02139, USA; $Department of Surgery, Harvard Medical School and Children’s Hospital, Boston, MA 02115, USA Polyglycolic incapable
acid (PGA) fibre of resisting
of poly(L-lactic
acid)
in chloroform PGA fibres
(PLLA)
were
The compression
bonded
tubes
degraded revealed
resisted
more
specific
structure
was exhibited
in vitro Tissue
formed
engineering,
PLGA.
Implantation
with
their
structure
by the finding a tubular
tissue
polyglycolic
with the extent
PLGA bonded
during lumen.
that smooth
acid, polylactic
dissolved
cells
Tubes
sprayed
tissue
by
on the
and PLLA bonded
bonded
of these
the
was controlled
of bonding,
of PLLA
fibrovascular
muscle
solutions
(PLGA)
of polymer
tubes.
The potential
with appropriate
acid) of bonding
mass
than
bonded
but they are atomized
The PLLA and PLGA coated
and extent
and the total
increased
forces
with a central
cells,
PGA meshes,
tubes.
The pattern
solution
of devices
compressive
maintained
into hollow
fibres.
in the atomized
to transplant
To stabilize
of poly(o,L-lactic-co-glycolic
formed
adjacent
than devices
of a tubular
tissues
onto devices Keywords:
slowly
candidates
forces.
copolymer
meshes
resistance larger
that the devices
the formation
over bonded
of polymer
device.
are attractive
compressional
and a 5060
sprayed
and physically
the concentration
meshes
significant
with
tubes
ingrowth, devices
and endothelial
PLLA
into rats
resulting
in
to engineer cells
seeded
cell distribution. acid,
smooth
muscle
cells,
endothelial
cells Received
26 October
1994; accepted
5 January
1995
to engineer a variety of tissues, including liver, cartilage and intestine3. This class of polymers degrades by a simple hydrolysis mechanism, and by varying the ratio of lactic and glycolic acids in the polymer one can control the crystallinity of the polymer, and thus its degradation rate and mechanical properties4. Furthermore, these polymers can be processed to yield a variety of different structures, including fibres, hollow tubes and porous sponges5-7. Non-woven meshes of polyglycolic acid (PGA) fibres have been particularly attractive materials for use as cell delivery devices as they are highly porous, permitting diffusion of nutrients throughout the device followimplantation ing while allowing subsequent neovascularization of the developing tissue, and they can be easily fabricated into devices with varying geometry. However, this material lacks the structural stability to withstand compressive forces in vivo, and external supports are necessary if one desires to form a stable three-dimensional structure (e.g. a tube) from this material” ‘. In this study, we investigated whether threedimensional structures capable of resisting large compressive forces and guiding the formation of a desired tissue structure could be formed from PGA fibre meshes by physically bonding adjacent fibres acid) using a spray casting method. Poly(L-lactic
While organ transplantation and tissue reconstruction are highly successful therapies for a variety of maladies, a shortage of donor tissue limits their application to a percentage of those who could potentially benefit from these therapies. For example, over 83 000 people either died or were maintained on less-thanoptimal therapies due to a lack of donated organs in the USA in 19901. To aid these people, a variety of investigators have proposed to engineer new tissues by transplanting isolated cell populations on biomaterial scaffolds to create functional new tissues in viva’. To engineer complex tissues such as blood vessels or intestine, cells must be localized to a specific site in vivo, and the formation of an appropriate tissue structure from the implanted cells and the host tissue must be promoted. Biodegradable materials are particularly attractive for fabricating the devices utilized to transplant cells and engineer new tissues because they can be designed to erode after tissue development is complete, leaving a completely natural tissue2’3. Templates synthesized from polymers of the lactic and glycolic acid family have previously been utilized ‘Current address: Departments of Biological and Materials Sciences and Chemical Engineering, University of Michigan, Ann Arbor, MI 48109, USA. Correspondence to Prof. R. Langer. 115
Biomaterials
1996. Vol. 17 No. 2
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116
(PLLA) or a 50/50 copolymer of lactic and glycolic acids (PLGA) was dissolved in chloroform, atomized and sprayed over a PGA mesh formed into a tubular structure. Following solvent evaporation, a physically bonded structure resulted, and the pattern and extent of PGA fibre bonding was controlled by the processing These tubular devices were capable of conditions. withstanding large compressive forces in vitro (SO200mN) and maintained their structure in vivo. The specific mechanical stability was dictated by the extent of physical bonding and the polymer utilized to bond the PGA fibres.
MATERIALS The PGA mesh (fibre diameter approximately 12 /lrn; mesh thickness = 0.3 mm, specific gravity = porosity = 97%) was purchased from 80.2 mg cm-“, Albany Int. (Taunton, MA, USA), PLLA and the Medisorb acid) from poly(D,L-lactic-co-glycolic (Cincinnati, OH, USA), the lactic dehydrogenase kit, glycolic and lactic acid standards, and 4,5-dihydroxy2,7_naphthalenedisodium salt were purchased from Sigma Chemical Co. (St Louis, MO, USA), chloroform from Mallinckrodt (Paris, KY, USA), phosphatebuffered saline and DMEM medium from Gibco (Grand Island, NY, USA), Tmax film from Kodak, Lewis rats (250-300g) from Charles River (Wilmington, MA, USA), calf serum from Hyclone Lab. Inc. (Logan, IJT, USA), penicillin and strepromycin from Irvine Scientific (Santa Ana, CA, USA), and methoxyflurane from Pitman-Moore Inc. (Mundelein, IL, USA).
METHODS Tube fabrication Rectangles (1.3 x 3.Ocm) of the non-woven mesh of PGA fibres were wrapped around a Teflon cylinder (outside diameter= 3.0mm) to form a tube, and the two overlapping ends were manually interlocked to form a seam. The Teflon cylinders were then rotated at 20rpm using a stirrer (Caframo; Wiarton, Ontario, Canada). Solutions of PLLA and PLGA dissolved in chloroform (l-15%, w/v) were placed in a dental atomizer (Devilbus Corp.) and sprayed over the rotating PGA mesh from a distance of 6in (~15 cm) using a nitrogen stream (18psi (-124.2 kPa)) to atomize the polymer solution. The PLGA and PLLA had molecular weights (M,) of 43400 (M,,,/i& = 1.43) and 74 100 (n/i,,./!& = 1.64), respectively. Molecular weights were determined by gel permeation chromatography as described previously7. While PLLA and copolymers of lactic and glycolic acids are soluble in PGA is very weakly soluble in this chloroform, solvent. Thus, the PGA fibres are largely unchanged by the process. After spraying was completed, the tubes were lyophilized to remove residual solvent, removed from the Teflon cylinder and cut into The tubes were sterilized by specific lengths. exposure to ethylene oxide for 24 h, followed by degassing for 24 h. Biomaterials
1996.
Vol.
17 No. 2
PGA tubes:
D.J. Mooney
et a/.
Device characterization The mass of PLLA and PLGA that bonded to the PGA scaffolds was determined by weighing PGA devices before and after spraying. For scanning electron microscopic examination, samples were gold coated using a Sputter Coater (Desk II, Denton Vacuum, Cherry Hill, NJ, USA). An environmental scanning electron microscope (ElectroScan, Wilmington, MA, USA) was operated at 30 kV with a water vapour environment of 5 torr (~665 Pa) to image samples. Photomicrographs were taken with Polaroid 55 film. Thermal mechanical analysis was performed with a TMA 7 (Perkin Elmer Corp, Norwalk, CT, USA) using a compression probe with a circular tip (d = 3.0 mm). All testing was done at a constant temperature of 37 ‘C. Tubes were placed on their sides for testing (axis of tube lumen perpendicular to the axis of force application), and the change in device diameter (parallel to the direction of force application) was followed during and after force application. The compressional forces applied to the tubes in vivo will presumably also be in a radial direction. The resulting deformations were normalized to the initial device diameter. Some samples were pre-wet by placing them in a vial containing phosphate-buffered saline and incubating at 37 ‘C for 24 h. All tests were performed in triplicate, and representative data are given. The erosion characteristics of bonded devices were assayed by placing individual tubes in 5ml of phosphate-buffered saline, pH 7.4, and incubating under static conditions at 37 C. The mass loss was analysed by weighing lyophilized devices before and after the incubation period. The release of lactic acid was assayed enzymatically with lactic dehydrogenase using a kit from Sigma. The release of glycolic acid was quantitiated with a calorimetric assay “’ which involves decarboxylating glycolic acid in the presence of concentrated sulphuric acid to form formaldehyde, followed by reaction of formaldehyde with chromotropic acid to yield a coloured product which can be quantitated spectrophotometrically.
Implantation of tubes Polymer constructs were implanted into the omentum of syngeneic Lewis rats as described previously”. NIH guidelines for the care and use of laboratory animals (NIH Publication No.85-23 Rev. 1985) have been observed in all experiments involving animals. Inhalation anaesthesia with methoxyflurane was always utilized. The omental tissue was rolled around the devices to promote tissue invasion and neovascularization of the implants from all sides. Implants were secured in place with sutures of 7-O Maxon (Davis and Geck). Recipients of polymer devices were killed on post-implantation days 3 and 18. The implants were removed, fixed in 10% buffered formalin and thin sections were cut from paraffin-embedded tissue. Histological sections were stained with haematoxylin and eosin. Photomicrographs were taken with Kodak Tmax film.
Cell seeding on devices To introduce (passage 6-9)
bovine aortic smooth muscle cells into the polymeric delivery devices,
Stabilized
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D.J.
Mooney et al.
1 ml of a cell suspension containing 5-20 x lo5 cellsmll’ was injected into the interior of each tube using a 1 ml syringe and a 22-gauge needle. The cell suspension was retained in the tubes by placing a small plug of the PGA fibres at both ends of the tubes during the cell adhesion period. Devices were incubated at 37°C in an atmosphere of 10% COZ to allow for cell adhesion and proliferation. The tubes were manually rotated periodically using sterile forceps during the period of cell adhesion to promote even cell seeding. Cell-polymer devices were kept in DMEM medium, containing 5% calf serum, 100 Umll’ penicillin and 100 mg mll’ streptomycin, during this time. The seeding protocol was repeated 7 days later to ensure even seeding of cells within the devices. Ten days later, a cell suspension of bovine aortic endothelial cells (passage 6-9) was similarly seeded onto the tubes. After 4 more days the devices were fixed in formalin, embedded in paraffin, sectioned and stained (haematoxylin and eosin) using standard techniques. Sections were stained for the presence of desmin (a smooth muscle specific protein) and Factor 8 (specific for endothelial cells) using standard immunohistochemical protocols. Antibodies for this analysis were purchased from Shandon (Pittsburgh, PA, USA). The endothelial cells and smooth muscle cells were isolated from bovine aortas using a collagenase digestion, and were a gift from Dr Judah Folkman.
RESULTS Bonding tubes with PLLA To determine whether PGA scaffolds could be stabilized by physically bonding adjacent fibres, chloroform containing dissolved PLLA (l-15% w/v) was sprayed over the exterior surface after the PGA mesh was wrapped around a Teflon cylinder to form a tube. The PLLA formed a coating over the exterior PGA fibres after the solvent evaporated, and physically bonded adjacent fibres. The tubes formed in this manner could be easily removed from the Teflon cylinder for characterization and use. The pattern of bonding was controlled by the concentration of the PLLA in the atomized solution (Figure I), even though the time of spraying was adjusted to maintain an approximately constant mass of PLLA on the devices under the various conditions (Table 1). Spraying with a solution containing 1 or 5% PLLA resulted in extensive bonding of PGA fibres without significantly blocking the pores of the PGA mesh. Spraying with a 10% solution of PLLA also bonded fibres, but resulted in the formation of a PLLA film on the exterior surface of the PGA mesh that contained only small pores. Spraying with a solution containing 15% PLLA had a similar effect, although the polymer film that formed was less organized. In all cases, the PLLA coated and bonded fibres only on the exterior surface of the PGA mesh, as no coating or bonding of fibres was observed on the interior surface of the PGA mesh (Figure 2). The compression resistance of bonded tubes was assessed in vitro to determine which patterns of
117 bonding resulted in the most stable devices. Unbonded tubes were completely crushed by a force of 5 mN, but bonded tubes were capable of resisting forces in excess of 200mN. However, the ability of bonded tubes to resist a given compressional force was dependent on the pattern of bonding (Figure 3). For example, tubes bonded with 1 or 15% PLLA were significantly compressed by a force of 200mN, while tubes bonded with a solution of 5 or 10% PLLA were only slightly compressed by this force. The compression was viscoelastic in all cases, as the devices only partially decompressed after the force was removed. Uniform properties were observed with respect to the position along and around a tube. To determine if the extent, as well as the pattern, of bonding could vary the compression resistance of tubes, an atomized dispersion of 5% PLLA was then sprayed over the devices for different times, Lengthening the spraying time from 10 to 60 s increased the mass of PLLA on the devices (Table 2). Infrequent bonds between adjacent fibres resulted from spraying for 10s. Spraying for more extended periods increased the PLLA coating over the PGA fibres, and the extent of bonding (Figure 4). The ability of these tubes to resist compressional forces and maintain their shape was quantitated again using thermal mechanical analysis. The compression resistance strongly depended on the extent of bonding, as tubes that were more extensively bonded had a greater resistance to deformation (Figure 5A). The compression that did occur under these conditions was again a combination of a reversible, elastic strain and an irreversible deformation. Some tubes were also exposed to an aqueous environment before testing to determine whether this environment for 24 h would destabilize the tubes. The aqueous environment had a slight detrimental effect on the stability of bonded tubes, but they were still capable of resisting large compressive forces (Figure 5B).
Bonding tubes with PLGA To determine whether this technique of stabilizing PGA devices could be utilized with a variety of polymers, the previous study was repeated using a 50150 copolymer of lactic and glycolic acids. The mass of polymer bonded to the devices and the extent of physical bonding were again regulated by the time an atomized dispersion of the bonding polymer was sprayed over the PGA fibres (Table 2; Figure 6). Once again, bonding increased the compression resistance of devices formed into a tubular structure (Figure ?‘A). However, these devices were not able to resist the same compressional forces as PLLA bonded devices. Tubes bonded with PLLA were capable of resisting forces up to 200mN, while tubes bonded with PLGA were only capable of resisting forces slightly greater than 5OmN. The difference between devices stabilized with PLLA and PLGA was even more striking when the devices were tested after immersion in phosphate-buffered saline for 24 h. PLGA bonded tubes, in contrast to PLLA bonded tubes, were significantly weakened by this treatment
(Figure 7B). Biomaterials
1996, Vol. 17
No. 2
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Fi gure 1 cc)ntaining of sprayed
D.J. Mooney et al.
Photomicrographs of the exterior surface of PGA meshes formed into tubular structures and sprayed (A) l%, (6) 5%, (C) 10% and (D) 15% PLLA. The spraying time was varied to yield an approximately PLLA in all conditions. The original magnifications and size bars are shown in the photomicrographs.
Table 1 PGA mesh concentration
sprayed
with
solutions
of varying
PLLA
PLLA concentration (w/v)
Spraying time (s)
Mass of PLLA on device* (X initial PGA mass)
1 5 10 15
150 30 15 10
115f20 168 i 16 145 Ik 12 108f73
‘Values
PGA tubes:
represent
the mea”
i
s.d.
of three
dewces
Tube degradation in vitro The time course for erosion of the tubes was determined by quantitating the mass loss and monomer release from tubes immersed in a pH balanced, isotonic saline solution. Devices bonded with PLGA were completely degraded by 11 weeks, while devices bonded with PLLA only lost 30% of their mass after 10 weeks (Figure 8A). The degradation of the PLLA bonded tubes was solely due to erosion of the PGA fibres, as glycolic acid was released from the Biomaterials 1996. Vol. 17 No. 2
with solution constant mas
S S
tubes, but virtually no lactic acid was released over this time from the tubes (Figure 8B). PLLA degrades and no significant loss of PLLA mass is slowly, expected until l-2 years. Erosion of tubes bonded with PLGA was due to erosion of both the PLGA fibres and the PLGA, as both glycolic acid and lactic acid were released from the tubes over this time frame (Figure SC).
Compression resistance in vivo To confirm that stabilized tubes were capable of resisting compressional forces in viva as well as in vitro, devices bonded with PLLA (5% PLLA; 30s spraying time) were implanted into the omentum of laboratory rats. The initial (3 day) host response was characterized by fibrin deposition and scattered inflammatory cells throughout the devices. A mature fibrovascular tissue was evident throughout the devices by 7 days, and the devices maintained their tubular structure with a central lumen for the 18 day duration of the experiment (Figure 9A). The invading fibroblasts and the newly deposited matrix were aligned with the lumens of the tubes (Figure 9B).
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Cell adhesion and organization in vitro on bonded tubes
Figure 2 A photomicrograph of the interior surface of PGA mesh formed into a tubular structure and sprayed with a solution of 5% PLLA for 30s. The interior surface, in contrast to the exterior surface (see Figure 7), was largely unaffected by this process. The original magnification and size bar are shown in the photomicrograph.
PLLA bonded tubes (5% PLLA; 30s spraying time) were subsequently seeded with smooth muscle cells and endothelial cells to investigate the suitability of these devices to serve as cell delivery vehicles. Blood vessels are largely comprised of these two cell types. The smooth muscle cells adhered to the polymer fibres (Figure ZOA and B), and proliferated to fill the void space present between polymer fibres (Figure ZUB). Endothelial cells also adhered to the devices, and over time formed a lining on the interior section of the devices (Figure IOA and C). lmmunohistochemical staining for desmin confirmed that the cells filling the interstices between polymer fibres were smooth muscle cells, and staining for Factor 8 confirmed that the cells lining the luminal surface were endothelial in nature (not shown). This organization of the muscle and endothelial cells is similar to that observed in blood vessels.
DISCUSSION
100 t
80
zz
5: ._
.-
I
I
10
Time
I
20
(min)
Figure 3 Representative strain diagrams of tubes formed from the PGA mesh after spraying with a solution containing 0, 1%; n , 5%; 0, 10%; and 0, 15% PLLA. Devices were subjected to a compressive force of 200 mN applied in a direction perpendicular to the axis of the device lumen starting at 0 min. The force was removed at 10 min, and the change in the diameter of the tube (parallel to the direction of force application) was monitored both during and after the time of force application, and normalized to the initial diameter.
Table 2 Spraying PGA scaffolds solution of PLLA or PLGA
for various
times
with a 5%
Spraying time (s)
Mass of PLLA on device’ (% initial PGA mass)
Mass of PLGA on device’ (% initial PGA mass)
10 20 30 60
43f 160 i 165 i 390 f
54 59 140 313
‘Values
represent
the
mean
11 55 22 37 i
s.d.
of three
devices.
i f f It
9 40 IO 51
Three-dimensional tubes can be formed from PGA fibre scaffolds by physically bonding adjacent fibres. The compression resistance and degradation rate of these devices were controlled by the pattern and extent of physical bonding, and the type of polymer utilized to bond the PGA fibres. Fibrovascular tissue invaded the leading to the devices following implantation, formation of a tubular tissue with a central lumen. The potential of these devices to engineer tissues was exhibited by the finding that endothelial cells and smooth muscle cells adhered to the devices and formed a new tissue in vitro with appropriate tissue organization. The compression resistance of devices was monitored by applying a constant force on the tubes. The resulting changes in the device diameters were partially elastic, as indicated by the partial decompression following removal of the applied force. The irreversible changes in the device diameters were likely caused by both crushing and bending of fibres, and by rearrangement of fibres. Contact between the compression tip and the tubes was not analysed, and will likely change as the tubes compress and fibres rearrange. For this reason, results were reported for compressional forces, not stresses. Calculation of stresses using the entire contact area of the compression probe would give the most conservative estimate of mechanical moduli. Tubes which were bonded with PLLA were more resistant to compressional forces than tubes bonded with PLGA. This finding is not surprising, as crystalline PLLA is typically much stiffer than amorphous PLGA4. Additionally, while the compression resistance of PLLA bonded devices was not greatly changed after exposure to an aqueous environment, PLGA bonded devices were markedly weakened after the same treatment. PLGA is more hydrophilic than PLLA4 due to the presence of the glycolic acid residues, and the absorbed water likely acts as a plasticizer, weakening Biomaterials
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Figure 4 containing crographs.
D.J.
Mooney et al.
Photomicrographs of the exterior surface of PGA meshes formed into tubular structures and sprayed with solutions 5% PLLA for (A) 10, (B) 20, (C) 30 and (D) 60s. The original magnifications and size bars are shown in the photomi-
the PLGA. The PLLA bonded devices were slightly weakened after this treatment, indicating that the PLLA was also somewhat plasticized. The erosion of the devices was also dependent on the polymer utilized for bonding. PLLA is hydrolysed very slowly, and virtually no lactic acid release was observed over the IO weeks of the erosion study. The erosion of PLLA bonded devices was entirely due to hydrolysis of the glycolic acid bonds in the fibres. In contrast, both the PGA fibres and the PLGA used to bond the fibres eroded completely over 11 weeks. The release of glycolic acid from these devices occurred more rapidly than the release of lactic acid. This was likely caused by the more rapid erosion of the PGA fibres, followed by the slower release of both lactic acid and glycolic acid from the PLGA. Biodegradable devices are attractive for cell transplantation and tissue engineering since they can be designed to erode once tissue development is complete, leaving a completely natural tissue. The approach described in this report to mechanically stabilize fibre-based scaffolds was performed with PGA, PLGA and PLLA because of the long history of these polymers in medical devices, and the range of degradation rates that can be obtained with this class Biomaterials
PGA tubes:
1996,
Vol. I 7 No. 2
of polymers (Figure 8). However, this technique could potentially be used with a variety of other polymers, both erodible and non-erodible, for medical or nonmedical applications. Various approaches have previously been taken to mechanically stabilize structures formed from PGA fibres. PGA fibres can be physically bonded with a second polymer in a similar manner as described here by simply dipping the PGA scaffold into a solution of PLLA dissolved in chloroform, and allowing the chloroform to evaporatel’. Alternatively, a thermal processing technique that results in temporary melting and subsequent bonding of PGA fibres has been reportedI”. The bonding approach described in this report is simple, permits a variety of bonding polymers to be utilized and allows the fabrication of various threedimensional scaffolds. It also results in bonding only of the outermost fibres of the device (Figure z), in contrast to the other methods. This preserves the desirable features of the PGA mesh (high porosity, high surface area/polymer mass ratio) throughout the interior sections. This approach also allows both the extent and pattern of bonding to be easily controlled. Extensive coating and bonding of fibres resulted when the polymer concentration in the atomized solution
Stabilized
PGA tubes:
D.J. Mooney
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121
60
set
30 20
set set
10
set
B Control 80
Pre-wet
; z’6 EZ ._ zz
60
ti 4
5.
40
z P
0 0
10
Time
(minutes)
20
0
5
Time
10
15
20
(minutes)
Figure 5 Representative strain diagrams of (A) PGA tubes sprayed for various times with a 5% PLLA solution and subjected to a compressive force of 200 mN starting at 0 min. The force was removed at 10 min. The force application and the change in the diameter of the tube (normalized to the initial diameter) were monitored, as described in the legend for Figure 3, both during and after the time of force application. (B) Devices sprayed with a 5% PLLA solution for 30s and tested dry (Control) or after pre-wetting for 24 h in a saline solution (Pre-wet). The compressional force was again 200 mN.
Figure 6
Photomicrographs of the exterior surface of PGA meshes formed into tubular structures and sprayed with solutions containing 5% PLGA for (A) 10, (B) 20, (C) 30 and (D) 60s. The original magnifications and size bars are shown in the photomicrographs. Biomaterials
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A
60
set
30
set
10
set
B
60
PGA tubes:
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::w Cont
loo*
Pre-wet
”
I
o-
,
10 (minutes)
0 Time
20
0
20
10
TIME
(minutes)
Figure 7 Representative strain diagrams of (A) PGA tubes sprayed for various times with a to a compressive force of 50 mN starting at 0 min. The force was removed at 10 minutes, and tube (normalized to the initial diameter) was monitored both during and after the time of force with a 5% PLGA solution for 30s and tested dry (Control) or after pre-wetting for 24 h in compressional force in this test was again 50 mN:
5% PLGA solution and subjected the change in the diameter of the application. (6) Devices sprayed a saline solution (Pre-wet). The
A 100
60
60
40
20
Lactic
acid
0 0
2
4
Time
6
6
10
12
0
2
4
(wk)
Time
Glycolic
60-
2
6
4
Time
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12
(wk)
6
6
acid
10
12
(wk)
Figure 8 The degradation of devices bonded by spraying with PLLA or PLGA (5% solution; spraying time measured by (A) quantitating the change in device mass over time, or (B) the release of glycolic and lactic acids bonded devices, or (C) PLGA bonded devices. Devices were incubated at 37’ C under static conditions in buffered removed at various times for analysis. Values in (A) represent the mean and standard deviation calculated samples. Biomaterials
10
acid
Lactic
0
6
=3Os), as from PLLA saline and from three
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Figure 9 (A) Low-power and (B) high-power photomicrographs of a histological section from a bonded tube (5% PLLA; 30s) implanted for 17 days in the omentum of a Lewis rat. These cross-sections of the implanted device were cut perpendicular to the axis of the tube’s lumen. (A) The central lumen (I) is visible, along with numerous polymer fibres (arrows), the host omental tissue (o), and the ingrown fibroblasts and fibrous tissue they deposited. (B) The fibroblasts which invaded the device and the fibrous tissue deposited by these cells aligned in parallel with the central lumen. The original magnifications of these photomicrographs were (A) x16 and (B) x158.
low (l-5%) (Figure IA and B). Increasing the concentration of polymer in the atomized solution to 10% resulted in the formation of a relatively smooth film over the external surface of PGA meshes, and utilizing a 15% solution resulted in the formation of a fibrous, non-homogeneous film over the PGA meshes (Figure ZC and D). Increasing the polymer concentration raises the viscosity of this solution and this likely increases the droplet size which is formed during the atomization process. This will effect how these droplets penetrate the PGA mesh, how they aggregate on the PGA mesh, and the rate of solvent evaporation. All of these factors will affect the pattern of bonding. To engineer a tissue with a desired three-dimensional structure, the cell delivery device must maintain a preconfigured geometry in the face of external forces during the process of tissue development. While the magnitude of the compressive forces that are exerted on implanted devices by the surrounding tissue are unclear, they are significant and will vary depending on the implant site. The magnitude of forces utilized in the present study to quantitate the compression resistance of devices in vitro was 50-2OOmN. This
was
Figure 10 (A) Low-power photomicrograph of a histological section of a bonded tube (5% PLLA; 30s) seeded with smooth muscle cells and endothelial cells in vitro as described in the Methods section. This cross-section was cut perpendicular to the axis of the tube’s lumen. Highpower photomicrographs of (B) an interior section of the device and (C) a section adjacent to the lumen. Smooth muscle cells readily adhered to polymer fibres (p) and filled the interstices between polymer fibres (A and B), while endothelial cells formed a lining on the luminal surface (A and C; arrows). The original magnifications of these photomicrographs were (A) x62.5, (B) x100 and (C) x158.
results in pressures ranging from approximately 50 to 200 mmHg (6.65-26.6 kPa) (assuming complete and continuous contact between the TMA compression tip and the tube). These pressures are in the same range Biomaterials
1996, Vol. 17 No. 2
Stabilized PGA tubes:D.J. Mooney
124 observed in blood vessels. Devices which were stable to high forces (PLLA bonded devices) were also stable after implantation into the omentum of laboratory rats. The omentum was chosen as the implant site because it is highly vascularized, easily accessed and manipulated and its anatomic location makes it a surgically, preferred site to engineer a variety of gastrointestinal tissues (e.g. small intestine). The compressional forces exerted by the surrounding tissue are likely not as great as other potential implant sites (e.g. popliteal space). The formation of fibrovascular tissue in implanted tubes was not surprising, as it is well documented that this type of ingrowth occurs in porous, synthetic materials'",14. The ingrowth and organization of the fibrovascular tissue will also exert compressional forces on the forming tissue, although the magnitude of these forces is unclear. It is anticipated that the ingrowing fibrovascular tissue would have eventually filled the central lumen of the implanted tubes since there was no epithelial cell lining of the lumen. An endothelial cell lining would likely prevent this outcome. While large diameter synthetic blood vessels (>5 mm diameter) have been successfully utilized for years, prosthetic small diameter blood vessels (~5 mm diameter) have been unsuccessful. Various investigators have attempted to improve the performance of small diameter grafts by either lining them with endothelial cells before implantation15, promoting endothelial cell migration from the adjacent vessels’“, or engineering blood vessels using extracellular matrix molecules as the template for tissue organization17. This last approach showed that the cell types utilized in the present study (endothelial and smooth muscle cells) have the ability to reform a tissue with the appropriate structure if placed on an appropriate matrix. The synthetic, biodegradable tubes described in this paper may provide a means to provide appropriate mechanical properties to an engineered blood vessel while also promoting the formation of a complete and natural replacement from the appropriate cell types. It may also be possible to combine the advantages of synthetic polymers (tailored mechanical and degradative properties, reproducible synthesis) with the biological specificity of extracellular matrix molecules such as collagen by producing templates from synthetic, biodegradable polymers which contain biologically active amino acid side chains’“. Using the appropriate cell signalling molecules on these polymers may allow one to promote endothelial cell adhesion in desirable spatial locations while preventing other cell types from adhering’“. The approach outlined in this report to stabilize fibre-based cell delivery devices could also be utilized to engineer a variety of other tubular tissues (e.g. intestine, trachea) and non-tubular tissues2”.
ef al.
REFERENCES 1
The
2
Sponsored by the United Network for Organ Sharing, Richmond, VA, USA. Langer R, Vacanti JP. Tissue engineering. Science 1993;
National
Cooperative
Transplantation
Study.
260:920-932. 3
Mooney DJ, Vacanti JP. Tissue engineering using cells and synthetic polymers. Transplant Rev 1993; 7: 153-
4
Gilding DK. Biodegradable polymers. In: Williams DF, ed. Biocompatibility of Clinical Implant Materials. Boca Raton, FL: CRC Press, 1981. Frazza EJ, Schmitt EE. A new absorbable suture. ]
162.
5
Biomed 6
7
8
Mater Res 1971;1:43.
Mooney DJ, Organ G, Vacanti JP, Langer R. Design and fabrication of cell delivery devices to engineer tubular tissues. Cell Transplant 1994; 3: 203-210. Mikos AG, Thorsen AJ, Czerwonka LA, Bao Y, Langer R. Preparation and characterization of poly(t.-lactic acid) foams. Polymer 1994; 35: 1068-1077. Organ GM, Mooney DJ, Hansen LK, Schloo B, Vacanti JP. Transplantation of enterocytes utilizing polymercell constructs to produce neointestine. Transplant
Proc 1992;24: 3008-3009. 9
Organ GM, Mooney DJ, Hansen LK, Schloo B, Vacanti JP. Enterocyte transplantation using cell-polymer devices to create intestinal cpithelial-lined tubes.
Transplant 10
Proc 1993; 25: 998-1001.
Tan S. A colorometric
assay for glycolic
acid. Clin Chim
Acta 1978: 89: 13-23. 11
12
Vacanti CA, Cima LG, Ratkowski D, Upton J, Vacanti JP. Tissue engineered growth of new cartilage in the shape of a human ear using synthetic polymers seeded with chondrocytes. In: Cima LG, Ron ES, eds. Tissue Inducing Biomaterials. Materials Research, Pittsburgh, PA, 1992; 252: 367-374. Mikos AG, Bao Y, Cima LG, Ingber DE, Vacanti JP, Langer R. Preparation of poly(glycolic acid) bonded fiber structures for cell attachment and transplantation.
J Biomed 33
14
Mater Res 1993; 27: 183.
Wesloski SA, Fries CC, Karlson KE, Bakey M, Sawyer PN. Porosity: primary determinant of ultimate fate of synthetic vascular grafts. Surgery 1961;50:91-96. White RA, Hirose FM, Sproat RW, Lawrence RS, Nelson RJ. Histopathologic observations after short-term implantation of porous elastomers in dogs. Biomater-
iak 1981; 2: 171-176. 15
Herring
MB. Endothelial
cell
seeding.
J Vascular
Surg
1991;5:731-732. 16
Clowes AW. Graft endothelialization: the role of angiogenic mechanisms. 1 Vascular Surg 1991; 5: 734-
17
Weinberg CB, Bell E. A blood from collagen and cultured
18
Barrera DA, Zylstra E, Lansbury PT, Langer R. Synthesis and RGD peptide modification of a new biodegradable copolymer: poly(lactic acid-co-lysine). J Am Chem Sot
19
Hubbell JA, Massia SP, Desai NP, Drumhellar PC. Endothelial cell-selective materials for tissue engineering in the vascular graft via a new receptor. Biotechnol-
20
Puelacher WC, Mooney D, Langer R, Upton J, Vacanti JP, Vacanti CA. Design of nasoseptal cartilage replacements synthesized from biodegradable polymers and chondrocytes. Biomaterials 1994;
736. vessel model constructed vascular cells. Science
1986;231:397-400.
1993; 115:11010-11011.
ACKNOWLEDGEMENTS The authors would like to acknowledge Dr Betsy Schloo for the preparation of histological sections. Financial support for this work was provided by the National Science Foundation (BCS-9202311) and Advanced Tissue Sciences.
Biomaterials
1996,Vol. 17 No. 2
ogy 1991; 9: 568-572.
15: 774-778.
ELSEVIER
Stabilized polyglycolic acid fibrebased tubes for tissue engineering D.J. Mooney *+$, C.L. Mazzoni*, C. Breued, K. McNamara*, D. Hem*, J.P. Vacanti’ and R. Langer* ‘Department of Chemical Engineering, Massachusetts institute of Technology, Cambridge, MA 02139, USA; $Department of Surgery, Harvard Medical School and Children’s Hospital, Boston, MA 02115, USA Polyglycolic incapable
acid (PGA) fibre of resisting
of poly(L-lactic
acid)
in chloroform PGA fibres
(PLLA)
were
The compression
bonded
tubes
degraded revealed
resisted
more
specific
structure
was exhibited
in vitro Tissue
formed
engineering,
PLGA.
Implantation
with
their
structure
by the finding a tubular
tissue
polyglycolic
with the extent
PLGA bonded
during lumen.
that smooth
acid, polylactic
dissolved
cells
Tubes
sprayed
tissue
by
on the
and PLLA bonded
bonded
of these
the
was controlled
of bonding,
of PLLA
fibrovascular
muscle
solutions
(PLGA)
of polymer
tubes.
The potential
with appropriate
acid) of bonding
mass
than
bonded
but they are atomized
The PLLA and PLGA coated
and extent
and the total
increased
forces
with a central
cells,
PGA meshes,
tubes.
The pattern
solution
of devices
compressive
maintained
into hollow
fibres.
in the atomized
to transplant
To stabilize
of poly(o,L-lactic-co-glycolic
formed
adjacent
than devices
of a tubular
tissues
onto devices Keywords:
slowly
candidates
forces.
copolymer
meshes
resistance larger
that the devices
the formation
over bonded
of polymer
device.
are attractive
compressional
and a 5060
sprayed
and physically
the concentration
meshes
significant
with
tubes
ingrowth, devices
and endothelial
PLLA
into rats
resulting
in
to engineer cells
seeded
cell distribution. acid,
smooth
muscle
cells,
endothelial
cells Received
26 October
1994; accepted
5 January
1995
to engineer a variety of tissues, including liver, cartilage and intestine3. This class of polymers degrades by a simple hydrolysis mechanism, and by varying the ratio of lactic and glycolic acids in the polymer one can control the crystallinity of the polymer, and thus its degradation rate and mechanical properties4. Furthermore, these polymers can be processed to yield a variety of different structures, including fibres, hollow tubes and porous sponges5-7. Non-woven meshes of polyglycolic acid (PGA) fibres have been particularly attractive materials for use as cell delivery devices as they are highly porous, permitting diffusion of nutrients throughout the device followimplantation ing while allowing subsequent neovascularization of the developing tissue, and they can be easily fabricated into devices with varying geometry. However, this material lacks the structural stability to withstand compressive forces in vivo, and external supports are necessary if one desires to form a stable three-dimensional structure (e.g. a tube) from this material” ‘. In this study, we investigated whether threedimensional structures capable of resisting large compressive forces and guiding the formation of a desired tissue structure could be formed from PGA fibre meshes by physically bonding adjacent fibres acid) using a spray casting method. Poly(L-lactic
While organ transplantation and tissue reconstruction are highly successful therapies for a variety of maladies, a shortage of donor tissue limits their application to a percentage of those who could potentially benefit from these therapies. For example, over 83 000 people either died or were maintained on less-thanoptimal therapies due to a lack of donated organs in the USA in 19901. To aid these people, a variety of investigators have proposed to engineer new tissues by transplanting isolated cell populations on biomaterial scaffolds to create functional new tissues in viva’. To engineer complex tissues such as blood vessels or intestine, cells must be localized to a specific site in vivo, and the formation of an appropriate tissue structure from the implanted cells and the host tissue must be promoted. Biodegradable materials are particularly attractive for fabricating the devices utilized to transplant cells and engineer new tissues because they can be designed to erode after tissue development is complete, leaving a completely natural tissue2’3. Templates synthesized from polymers of the lactic and glycolic acid family have previously been utilized ‘Current address: Departments of Biological and Materials Sciences and Chemical Engineering, University of Michigan, Ann Arbor, MI 48109, USA. Correspondence to Prof. R. Langer. 115
Biomaterials
1996. Vol. 17 No. 2
Stabilized
116
(PLLA) or a 50/50 copolymer of lactic and glycolic acids (PLGA) was dissolved in chloroform, atomized and sprayed over a PGA mesh formed into a tubular structure. Following solvent evaporation, a physically bonded structure resulted, and the pattern and extent of PGA fibre bonding was controlled by the processing These tubular devices were capable of conditions. withstanding large compressive forces in vitro (SO200mN) and maintained their structure in vivo. The specific mechanical stability was dictated by the extent of physical bonding and the polymer utilized to bond the PGA fibres.
MATERIALS The PGA mesh (fibre diameter approximately 12 /lrn; mesh thickness = 0.3 mm, specific gravity = porosity = 97%) was purchased from 80.2 mg cm-“, Albany Int. (Taunton, MA, USA), PLLA and the Medisorb acid) from poly(D,L-lactic-co-glycolic (Cincinnati, OH, USA), the lactic dehydrogenase kit, glycolic and lactic acid standards, and 4,5-dihydroxy2,7_naphthalenedisodium salt were purchased from Sigma Chemical Co. (St Louis, MO, USA), chloroform from Mallinckrodt (Paris, KY, USA), phosphatebuffered saline and DMEM medium from Gibco (Grand Island, NY, USA), Tmax film from Kodak, Lewis rats (250-300g) from Charles River (Wilmington, MA, USA), calf serum from Hyclone Lab. Inc. (Logan, IJT, USA), penicillin and strepromycin from Irvine Scientific (Santa Ana, CA, USA), and methoxyflurane from Pitman-Moore Inc. (Mundelein, IL, USA).
METHODS Tube fabrication Rectangles (1.3 x 3.Ocm) of the non-woven mesh of PGA fibres were wrapped around a Teflon cylinder (outside diameter= 3.0mm) to form a tube, and the two overlapping ends were manually interlocked to form a seam. The Teflon cylinders were then rotated at 20rpm using a stirrer (Caframo; Wiarton, Ontario, Canada). Solutions of PLLA and PLGA dissolved in chloroform (l-15%, w/v) were placed in a dental atomizer (Devilbus Corp.) and sprayed over the rotating PGA mesh from a distance of 6in (~15 cm) using a nitrogen stream (18psi (-124.2 kPa)) to atomize the polymer solution. The PLGA and PLLA had molecular weights (M,) of 43400 (M,,,/i& = 1.43) and 74 100 (n/i,,./!& = 1.64), respectively. Molecular weights were determined by gel permeation chromatography as described previously7. While PLLA and copolymers of lactic and glycolic acids are soluble in PGA is very weakly soluble in this chloroform, solvent. Thus, the PGA fibres are largely unchanged by the process. After spraying was completed, the tubes were lyophilized to remove residual solvent, removed from the Teflon cylinder and cut into The tubes were sterilized by specific lengths. exposure to ethylene oxide for 24 h, followed by degassing for 24 h. Biomaterials
1996.
Vol.
17 No. 2
PGA tubes:
D.J. Mooney
et a/.
Device characterization The mass of PLLA and PLGA that bonded to the PGA scaffolds was determined by weighing PGA devices before and after spraying. For scanning electron microscopic examination, samples were gold coated using a Sputter Coater (Desk II, Denton Vacuum, Cherry Hill, NJ, USA). An environmental scanning electron microscope (ElectroScan, Wilmington, MA, USA) was operated at 30 kV with a water vapour environment of 5 torr (~665 Pa) to image samples. Photomicrographs were taken with Polaroid 55 film. Thermal mechanical analysis was performed with a TMA 7 (Perkin Elmer Corp, Norwalk, CT, USA) using a compression probe with a circular tip (d = 3.0 mm). All testing was done at a constant temperature of 37 ‘C. Tubes were placed on their sides for testing (axis of tube lumen perpendicular to the axis of force application), and the change in device diameter (parallel to the direction of force application) was followed during and after force application. The compressional forces applied to the tubes in vivo will presumably also be in a radial direction. The resulting deformations were normalized to the initial device diameter. Some samples were pre-wet by placing them in a vial containing phosphate-buffered saline and incubating at 37 ‘C for 24 h. All tests were performed in triplicate, and representative data are given. The erosion characteristics of bonded devices were assayed by placing individual tubes in 5ml of phosphate-buffered saline, pH 7.4, and incubating under static conditions at 37 C. The mass loss was analysed by weighing lyophilized devices before and after the incubation period. The release of lactic acid was assayed enzymatically with lactic dehydrogenase using a kit from Sigma. The release of glycolic acid was quantitiated with a calorimetric assay “’ which involves decarboxylating glycolic acid in the presence of concentrated sulphuric acid to form formaldehyde, followed by reaction of formaldehyde with chromotropic acid to yield a coloured product which can be quantitated spectrophotometrically.
Implantation of tubes Polymer constructs were implanted into the omentum of syngeneic Lewis rats as described previously”. NIH guidelines for the care and use of laboratory animals (NIH Publication No.85-23 Rev. 1985) have been observed in all experiments involving animals. Inhalation anaesthesia with methoxyflurane was always utilized. The omental tissue was rolled around the devices to promote tissue invasion and neovascularization of the implants from all sides. Implants were secured in place with sutures of 7-O Maxon (Davis and Geck). Recipients of polymer devices were killed on post-implantation days 3 and 18. The implants were removed, fixed in 10% buffered formalin and thin sections were cut from paraffin-embedded tissue. Histological sections were stained with haematoxylin and eosin. Photomicrographs were taken with Kodak Tmax film.
Cell seeding on devices To introduce (passage 6-9)
bovine aortic smooth muscle cells into the polymeric delivery devices,
Stabilized
PGA tubes:
D.J.
Mooney et al.
1 ml of a cell suspension containing 5-20 x lo5 cellsmll’ was injected into the interior of each tube using a 1 ml syringe and a 22-gauge needle. The cell suspension was retained in the tubes by placing a small plug of the PGA fibres at both ends of the tubes during the cell adhesion period. Devices were incubated at 37°C in an atmosphere of 10% COZ to allow for cell adhesion and proliferation. The tubes were manually rotated periodically using sterile forceps during the period of cell adhesion to promote even cell seeding. Cell-polymer devices were kept in DMEM medium, containing 5% calf serum, 100 Umll’ penicillin and 100 mg mll’ streptomycin, during this time. The seeding protocol was repeated 7 days later to ensure even seeding of cells within the devices. Ten days later, a cell suspension of bovine aortic endothelial cells (passage 6-9) was similarly seeded onto the tubes. After 4 more days the devices were fixed in formalin, embedded in paraffin, sectioned and stained (haematoxylin and eosin) using standard techniques. Sections were stained for the presence of desmin (a smooth muscle specific protein) and Factor 8 (specific for endothelial cells) using standard immunohistochemical protocols. Antibodies for this analysis were purchased from Shandon (Pittsburgh, PA, USA). The endothelial cells and smooth muscle cells were isolated from bovine aortas using a collagenase digestion, and were a gift from Dr Judah Folkman.
RESULTS Bonding tubes with PLLA To determine whether PGA scaffolds could be stabilized by physically bonding adjacent fibres, chloroform containing dissolved PLLA (l-15% w/v) was sprayed over the exterior surface after the PGA mesh was wrapped around a Teflon cylinder to form a tube. The PLLA formed a coating over the exterior PGA fibres after the solvent evaporated, and physically bonded adjacent fibres. The tubes formed in this manner could be easily removed from the Teflon cylinder for characterization and use. The pattern of bonding was controlled by the concentration of the PLLA in the atomized solution (Figure I), even though the time of spraying was adjusted to maintain an approximately constant mass of PLLA on the devices under the various conditions (Table 1). Spraying with a solution containing 1 or 5% PLLA resulted in extensive bonding of PGA fibres without significantly blocking the pores of the PGA mesh. Spraying with a 10% solution of PLLA also bonded fibres, but resulted in the formation of a PLLA film on the exterior surface of the PGA mesh that contained only small pores. Spraying with a solution containing 15% PLLA had a similar effect, although the polymer film that formed was less organized. In all cases, the PLLA coated and bonded fibres only on the exterior surface of the PGA mesh, as no coating or bonding of fibres was observed on the interior surface of the PGA mesh (Figure 2). The compression resistance of bonded tubes was assessed in vitro to determine which patterns of
117 bonding resulted in the most stable devices. Unbonded tubes were completely crushed by a force of 5 mN, but bonded tubes were capable of resisting forces in excess of 200mN. However, the ability of bonded tubes to resist a given compressional force was dependent on the pattern of bonding (Figure 3). For example, tubes bonded with 1 or 15% PLLA were significantly compressed by a force of 200mN, while tubes bonded with a solution of 5 or 10% PLLA were only slightly compressed by this force. The compression was viscoelastic in all cases, as the devices only partially decompressed after the force was removed. Uniform properties were observed with respect to the position along and around a tube. To determine if the extent, as well as the pattern, of bonding could vary the compression resistance of tubes, an atomized dispersion of 5% PLLA was then sprayed over the devices for different times, Lengthening the spraying time from 10 to 60 s increased the mass of PLLA on the devices (Table 2). Infrequent bonds between adjacent fibres resulted from spraying for 10s. Spraying for more extended periods increased the PLLA coating over the PGA fibres, and the extent of bonding (Figure 4). The ability of these tubes to resist compressional forces and maintain their shape was quantitated again using thermal mechanical analysis. The compression resistance strongly depended on the extent of bonding, as tubes that were more extensively bonded had a greater resistance to deformation (Figure 5A). The compression that did occur under these conditions was again a combination of a reversible, elastic strain and an irreversible deformation. Some tubes were also exposed to an aqueous environment before testing to determine whether this environment for 24 h would destabilize the tubes. The aqueous environment had a slight detrimental effect on the stability of bonded tubes, but they were still capable of resisting large compressive forces (Figure 5B).
Bonding tubes with PLGA To determine whether this technique of stabilizing PGA devices could be utilized with a variety of polymers, the previous study was repeated using a 50150 copolymer of lactic and glycolic acids. The mass of polymer bonded to the devices and the extent of physical bonding were again regulated by the time an atomized dispersion of the bonding polymer was sprayed over the PGA fibres (Table 2; Figure 6). Once again, bonding increased the compression resistance of devices formed into a tubular structure (Figure ?‘A). However, these devices were not able to resist the same compressional forces as PLLA bonded devices. Tubes bonded with PLLA were capable of resisting forces up to 200mN, while tubes bonded with PLGA were only capable of resisting forces slightly greater than 5OmN. The difference between devices stabilized with PLLA and PLGA was even more striking when the devices were tested after immersion in phosphate-buffered saline for 24 h. PLGA bonded tubes, in contrast to PLLA bonded tubes, were significantly weakened by this treatment
(Figure 7B). Biomaterials
1996, Vol. 17
No. 2
Stabilized
118
Fi gure 1 cc)ntaining of sprayed
D.J. Mooney et al.
Photomicrographs of the exterior surface of PGA meshes formed into tubular structures and sprayed (A) l%, (6) 5%, (C) 10% and (D) 15% PLLA. The spraying time was varied to yield an approximately PLLA in all conditions. The original magnifications and size bars are shown in the photomicrographs.
Table 1 PGA mesh concentration
sprayed
with
solutions
of varying
PLLA
PLLA concentration (w/v)
Spraying time (s)
Mass of PLLA on device* (X initial PGA mass)
1 5 10 15
150 30 15 10
115f20 168 i 16 145 Ik 12 108f73
‘Values
PGA tubes:
represent
the mea”
i
s.d.
of three
dewces
Tube degradation in vitro The time course for erosion of the tubes was determined by quantitating the mass loss and monomer release from tubes immersed in a pH balanced, isotonic saline solution. Devices bonded with PLGA were completely degraded by 11 weeks, while devices bonded with PLLA only lost 30% of their mass after 10 weeks (Figure 8A). The degradation of the PLLA bonded tubes was solely due to erosion of the PGA fibres, as glycolic acid was released from the Biomaterials 1996. Vol. 17 No. 2
with solution constant mas
S S
tubes, but virtually no lactic acid was released over this time from the tubes (Figure 8B). PLLA degrades and no significant loss of PLLA mass is slowly, expected until l-2 years. Erosion of tubes bonded with PLGA was due to erosion of both the PLGA fibres and the PLGA, as both glycolic acid and lactic acid were released from the tubes over this time frame (Figure SC).
Compression resistance in vivo To confirm that stabilized tubes were capable of resisting compressional forces in viva as well as in vitro, devices bonded with PLLA (5% PLLA; 30s spraying time) were implanted into the omentum of laboratory rats. The initial (3 day) host response was characterized by fibrin deposition and scattered inflammatory cells throughout the devices. A mature fibrovascular tissue was evident throughout the devices by 7 days, and the devices maintained their tubular structure with a central lumen for the 18 day duration of the experiment (Figure 9A). The invading fibroblasts and the newly deposited matrix were aligned with the lumens of the tubes (Figure 9B).
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D.J. Mooney et al.
119
Cell adhesion and organization in vitro on bonded tubes
Figure 2 A photomicrograph of the interior surface of PGA mesh formed into a tubular structure and sprayed with a solution of 5% PLLA for 30s. The interior surface, in contrast to the exterior surface (see Figure 7), was largely unaffected by this process. The original magnification and size bar are shown in the photomicrograph.
PLLA bonded tubes (5% PLLA; 30s spraying time) were subsequently seeded with smooth muscle cells and endothelial cells to investigate the suitability of these devices to serve as cell delivery vehicles. Blood vessels are largely comprised of these two cell types. The smooth muscle cells adhered to the polymer fibres (Figure ZOA and B), and proliferated to fill the void space present between polymer fibres (Figure ZUB). Endothelial cells also adhered to the devices, and over time formed a lining on the interior section of the devices (Figure IOA and C). lmmunohistochemical staining for desmin confirmed that the cells filling the interstices between polymer fibres were smooth muscle cells, and staining for Factor 8 confirmed that the cells lining the luminal surface were endothelial in nature (not shown). This organization of the muscle and endothelial cells is similar to that observed in blood vessels.
DISCUSSION
100 t
80
zz
5: ._
.-
I
I
10
Time
I
20
(min)
Figure 3 Representative strain diagrams of tubes formed from the PGA mesh after spraying with a solution containing 0, 1%; n , 5%; 0, 10%; and 0, 15% PLLA. Devices were subjected to a compressive force of 200 mN applied in a direction perpendicular to the axis of the device lumen starting at 0 min. The force was removed at 10 min, and the change in the diameter of the tube (parallel to the direction of force application) was monitored both during and after the time of force application, and normalized to the initial diameter.
Table 2 Spraying PGA scaffolds solution of PLLA or PLGA
for various
times
with a 5%
Spraying time (s)
Mass of PLLA on device’ (% initial PGA mass)
Mass of PLGA on device’ (% initial PGA mass)
10 20 30 60
43f 160 i 165 i 390 f
54 59 140 313
‘Values
represent
the
mean
11 55 22 37 i
s.d.
of three
devices.
i f f It
9 40 IO 51
Three-dimensional tubes can be formed from PGA fibre scaffolds by physically bonding adjacent fibres. The compression resistance and degradation rate of these devices were controlled by the pattern and extent of physical bonding, and the type of polymer utilized to bond the PGA fibres. Fibrovascular tissue invaded the leading to the devices following implantation, formation of a tubular tissue with a central lumen. The potential of these devices to engineer tissues was exhibited by the finding that endothelial cells and smooth muscle cells adhered to the devices and formed a new tissue in vitro with appropriate tissue organization. The compression resistance of devices was monitored by applying a constant force on the tubes. The resulting changes in the device diameters were partially elastic, as indicated by the partial decompression following removal of the applied force. The irreversible changes in the device diameters were likely caused by both crushing and bending of fibres, and by rearrangement of fibres. Contact between the compression tip and the tubes was not analysed, and will likely change as the tubes compress and fibres rearrange. For this reason, results were reported for compressional forces, not stresses. Calculation of stresses using the entire contact area of the compression probe would give the most conservative estimate of mechanical moduli. Tubes which were bonded with PLLA were more resistant to compressional forces than tubes bonded with PLGA. This finding is not surprising, as crystalline PLLA is typically much stiffer than amorphous PLGA4. Additionally, while the compression resistance of PLLA bonded devices was not greatly changed after exposure to an aqueous environment, PLGA bonded devices were markedly weakened after the same treatment. PLGA is more hydrophilic than PLLA4 due to the presence of the glycolic acid residues, and the absorbed water likely acts as a plasticizer, weakening Biomaterials
1996, Vol. 17 No. 2
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Figure 4 containing crographs.
D.J.
Mooney et al.
Photomicrographs of the exterior surface of PGA meshes formed into tubular structures and sprayed with solutions 5% PLLA for (A) 10, (B) 20, (C) 30 and (D) 60s. The original magnifications and size bars are shown in the photomi-
the PLGA. The PLLA bonded devices were slightly weakened after this treatment, indicating that the PLLA was also somewhat plasticized. The erosion of the devices was also dependent on the polymer utilized for bonding. PLLA is hydrolysed very slowly, and virtually no lactic acid release was observed over the IO weeks of the erosion study. The erosion of PLLA bonded devices was entirely due to hydrolysis of the glycolic acid bonds in the fibres. In contrast, both the PGA fibres and the PLGA used to bond the fibres eroded completely over 11 weeks. The release of glycolic acid from these devices occurred more rapidly than the release of lactic acid. This was likely caused by the more rapid erosion of the PGA fibres, followed by the slower release of both lactic acid and glycolic acid from the PLGA. Biodegradable devices are attractive for cell transplantation and tissue engineering since they can be designed to erode once tissue development is complete, leaving a completely natural tissue. The approach described in this report to mechanically stabilize fibre-based scaffolds was performed with PGA, PLGA and PLLA because of the long history of these polymers in medical devices, and the range of degradation rates that can be obtained with this class Biomaterials
PGA tubes:
1996,
Vol. I 7 No. 2
of polymers (Figure 8). However, this technique could potentially be used with a variety of other polymers, both erodible and non-erodible, for medical or nonmedical applications. Various approaches have previously been taken to mechanically stabilize structures formed from PGA fibres. PGA fibres can be physically bonded with a second polymer in a similar manner as described here by simply dipping the PGA scaffold into a solution of PLLA dissolved in chloroform, and allowing the chloroform to evaporatel’. Alternatively, a thermal processing technique that results in temporary melting and subsequent bonding of PGA fibres has been reportedI”. The bonding approach described in this report is simple, permits a variety of bonding polymers to be utilized and allows the fabrication of various threedimensional scaffolds. It also results in bonding only of the outermost fibres of the device (Figure z), in contrast to the other methods. This preserves the desirable features of the PGA mesh (high porosity, high surface area/polymer mass ratio) throughout the interior sections. This approach also allows both the extent and pattern of bonding to be easily controlled. Extensive coating and bonding of fibres resulted when the polymer concentration in the atomized solution
Stabilized
PGA tubes:
D.J. Mooney
et al.
121
60
set
30 20
set set
10
set
B Control 80
Pre-wet
; z’6 EZ ._ zz
60
ti 4
5.
40
z P
0 0
10
Time
(minutes)
20
0
5
Time
10
15
20
(minutes)
Figure 5 Representative strain diagrams of (A) PGA tubes sprayed for various times with a 5% PLLA solution and subjected to a compressive force of 200 mN starting at 0 min. The force was removed at 10 min. The force application and the change in the diameter of the tube (normalized to the initial diameter) were monitored, as described in the legend for Figure 3, both during and after the time of force application. (B) Devices sprayed with a 5% PLLA solution for 30s and tested dry (Control) or after pre-wetting for 24 h in a saline solution (Pre-wet). The compressional force was again 200 mN.
Figure 6
Photomicrographs of the exterior surface of PGA meshes formed into tubular structures and sprayed with solutions containing 5% PLGA for (A) 10, (B) 20, (C) 30 and (D) 60s. The original magnifications and size bars are shown in the photomicrographs. Biomaterials
1996, Vol.
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2
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A
60
set
30
set
10
set
B
60
PGA tubes:
D.J. Moonev
et al.
::w Cont
loo*
Pre-wet
”
I
o-
,
10 (minutes)
0 Time
20
0
20
10
TIME
(minutes)
Figure 7 Representative strain diagrams of (A) PGA tubes sprayed for various times with a to a compressive force of 50 mN starting at 0 min. The force was removed at 10 minutes, and tube (normalized to the initial diameter) was monitored both during and after the time of force with a 5% PLGA solution for 30s and tested dry (Control) or after pre-wetting for 24 h in compressional force in this test was again 50 mN:
5% PLGA solution and subjected the change in the diameter of the application. (6) Devices sprayed a saline solution (Pre-wet). The
A 100
60
60
40
20
Lactic
acid
0 0
2
4
Time
6
6
10
12
0
2
4
(wk)
Time
Glycolic
60-
2
6
4
Time
1996,
Vol.
17 No. 2
12
(wk)
6
6
acid
10
12
(wk)
Figure 8 The degradation of devices bonded by spraying with PLLA or PLGA (5% solution; spraying time measured by (A) quantitating the change in device mass over time, or (B) the release of glycolic and lactic acids bonded devices, or (C) PLGA bonded devices. Devices were incubated at 37’ C under static conditions in buffered removed at various times for analysis. Values in (A) represent the mean and standard deviation calculated samples. Biomaterials
10
acid
Lactic
0
6
=3Os), as from PLLA saline and from three
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123
Figure 9 (A) Low-power and (B) high-power photomicrographs of a histological section from a bonded tube (5% PLLA; 30s) implanted for 17 days in the omentum of a Lewis rat. These cross-sections of the implanted device were cut perpendicular to the axis of the tube’s lumen. (A) The central lumen (I) is visible, along with numerous polymer fibres (arrows), the host omental tissue (o), and the ingrown fibroblasts and fibrous tissue they deposited. (B) The fibroblasts which invaded the device and the fibrous tissue deposited by these cells aligned in parallel with the central lumen. The original magnifications of these photomicrographs were (A) x16 and (B) x158.
low (l-5%) (Figure IA and B). Increasing the concentration of polymer in the atomized solution to 10% resulted in the formation of a relatively smooth film over the external surface of PGA meshes, and utilizing a 15% solution resulted in the formation of a fibrous, non-homogeneous film over the PGA meshes (Figure ZC and D). Increasing the polymer concentration raises the viscosity of this solution and this likely increases the droplet size which is formed during the atomization process. This will effect how these droplets penetrate the PGA mesh, how they aggregate on the PGA mesh, and the rate of solvent evaporation. All of these factors will affect the pattern of bonding. To engineer a tissue with a desired three-dimensional structure, the cell delivery device must maintain a preconfigured geometry in the face of external forces during the process of tissue development. While the magnitude of the compressive forces that are exerted on implanted devices by the surrounding tissue are unclear, they are significant and will vary depending on the implant site. The magnitude of forces utilized in the present study to quantitate the compression resistance of devices in vitro was 50-2OOmN. This
was
Figure 10 (A) Low-power photomicrograph of a histological section of a bonded tube (5% PLLA; 30s) seeded with smooth muscle cells and endothelial cells in vitro as described in the Methods section. This cross-section was cut perpendicular to the axis of the tube’s lumen. Highpower photomicrographs of (B) an interior section of the device and (C) a section adjacent to the lumen. Smooth muscle cells readily adhered to polymer fibres (p) and filled the interstices between polymer fibres (A and B), while endothelial cells formed a lining on the luminal surface (A and C; arrows). The original magnifications of these photomicrographs were (A) x62.5, (B) x100 and (C) x158.
results in pressures ranging from approximately 50 to 200 mmHg (6.65-26.6 kPa) (assuming complete and continuous contact between the TMA compression tip and the tube). These pressures are in the same range Biomaterials
1996, Vol. 17 No. 2
Stabilized PGA tubes:D.J. Mooney
124 observed in blood vessels. Devices which were stable to high forces (PLLA bonded devices) were also stable after implantation into the omentum of laboratory rats. The omentum was chosen as the implant site because it is highly vascularized, easily accessed and manipulated and its anatomic location makes it a surgically, preferred site to engineer a variety of gastrointestinal tissues (e.g. small intestine). The compressional forces exerted by the surrounding tissue are likely not as great as other potential implant sites (e.g. popliteal space). The formation of fibrovascular tissue in implanted tubes was not surprising, as it is well documented that this type of ingrowth occurs in porous, synthetic materials'",14. The ingrowth and organization of the fibrovascular tissue will also exert compressional forces on the forming tissue, although the magnitude of these forces is unclear. It is anticipated that the ingrowing fibrovascular tissue would have eventually filled the central lumen of the implanted tubes since there was no epithelial cell lining of the lumen. An endothelial cell lining would likely prevent this outcome. While large diameter synthetic blood vessels (>5 mm diameter) have been successfully utilized for years, prosthetic small diameter blood vessels (~5 mm diameter) have been unsuccessful. Various investigators have attempted to improve the performance of small diameter grafts by either lining them with endothelial cells before implantation15, promoting endothelial cell migration from the adjacent vessels’“, or engineering blood vessels using extracellular matrix molecules as the template for tissue organization17. This last approach showed that the cell types utilized in the present study (endothelial and smooth muscle cells) have the ability to reform a tissue with the appropriate structure if placed on an appropriate matrix. The synthetic, biodegradable tubes described in this paper may provide a means to provide appropriate mechanical properties to an engineered blood vessel while also promoting the formation of a complete and natural replacement from the appropriate cell types. It may also be possible to combine the advantages of synthetic polymers (tailored mechanical and degradative properties, reproducible synthesis) with the biological specificity of extracellular matrix molecules such as collagen by producing templates from synthetic, biodegradable polymers which contain biologically active amino acid side chains’“. Using the appropriate cell signalling molecules on these polymers may allow one to promote endothelial cell adhesion in desirable spatial locations while preventing other cell types from adhering’“. The approach outlined in this report to stabilize fibre-based cell delivery devices could also be utilized to engineer a variety of other tubular tissues (e.g. intestine, trachea) and non-tubular tissues2”.
ef al.
REFERENCES 1
The
2
Sponsored by the United Network for Organ Sharing, Richmond, VA, USA. Langer R, Vacanti JP. Tissue engineering. Science 1993;
National
Cooperative
Transplantation
Study.
260:920-932. 3
Mooney DJ, Vacanti JP. Tissue engineering using cells and synthetic polymers. Transplant Rev 1993; 7: 153-
4
Gilding DK. Biodegradable polymers. In: Williams DF, ed. Biocompatibility of Clinical Implant Materials. Boca Raton, FL: CRC Press, 1981. Frazza EJ, Schmitt EE. A new absorbable suture. ]
162.
5
Biomed 6
7
8
Mater Res 1971;1:43.
Mooney DJ, Organ G, Vacanti JP, Langer R. Design and fabrication of cell delivery devices to engineer tubular tissues. Cell Transplant 1994; 3: 203-210. Mikos AG, Thorsen AJ, Czerwonka LA, Bao Y, Langer R. Preparation and characterization of poly(t.-lactic acid) foams. Polymer 1994; 35: 1068-1077. Organ GM, Mooney DJ, Hansen LK, Schloo B, Vacanti JP. Transplantation of enterocytes utilizing polymercell constructs to produce neointestine. Transplant
Proc 1992;24: 3008-3009. 9
Organ GM, Mooney DJ, Hansen LK, Schloo B, Vacanti JP. Enterocyte transplantation using cell-polymer devices to create intestinal cpithelial-lined tubes.
Transplant 10
Proc 1993; 25: 998-1001.
Tan S. A colorometric
assay for glycolic
acid. Clin Chim
Acta 1978: 89: 13-23. 11
12
Vacanti CA, Cima LG, Ratkowski D, Upton J, Vacanti JP. Tissue engineered growth of new cartilage in the shape of a human ear using synthetic polymers seeded with chondrocytes. In: Cima LG, Ron ES, eds. Tissue Inducing Biomaterials. Materials Research, Pittsburgh, PA, 1992; 252: 367-374. Mikos AG, Bao Y, Cima LG, Ingber DE, Vacanti JP, Langer R. Preparation of poly(glycolic acid) bonded fiber structures for cell attachment and transplantation.
J Biomed 33
14
Mater Res 1993; 27: 183.
Wesloski SA, Fries CC, Karlson KE, Bakey M, Sawyer PN. Porosity: primary determinant of ultimate fate of synthetic vascular grafts. Surgery 1961;50:91-96. White RA, Hirose FM, Sproat RW, Lawrence RS, Nelson RJ. Histopathologic observations after short-term implantation of porous elastomers in dogs. Biomater-
iak 1981; 2: 171-176. 15
Herring
MB. Endothelial
cell
seeding.
J Vascular
Surg
1991;5:731-732. 16
Clowes AW. Graft endothelialization: the role of angiogenic mechanisms. 1 Vascular Surg 1991; 5: 734-
17
Weinberg CB, Bell E. A blood from collagen and cultured
18
Barrera DA, Zylstra E, Lansbury PT, Langer R. Synthesis and RGD peptide modification of a new biodegradable copolymer: poly(lactic acid-co-lysine). J Am Chem Sot
19
Hubbell JA, Massia SP, Desai NP, Drumhellar PC. Endothelial cell-selective materials for tissue engineering in the vascular graft via a new receptor. Biotechnol-
20
Puelacher WC, Mooney D, Langer R, Upton J, Vacanti JP, Vacanti CA. Design of nasoseptal cartilage replacements synthesized from biodegradable polymers and chondrocytes. Biomaterials 1994;
736. vessel model constructed vascular cells. Science
1986;231:397-400.
1993; 115:11010-11011.
ACKNOWLEDGEMENTS The authors would like to acknowledge Dr Betsy Schloo for the preparation of histological sections. Financial support for this work was provided by the National Science Foundation (BCS-9202311) and Advanced Tissue Sciences.
Biomaterials
1996,Vol. 17 No. 2
ogy 1991; 9: 568-572.
15: 774-778.