Cardiovascular Research 63 (2004) 719 – 730
www.elsevier.com/locate/cardiores
Tissue engineering of aortic tissue: dire consequence of suboptimal elastic
fiber synthesis in vivo
Florian Opitz a,*, Katja Schenke-Layland a, Tina U. Cohnert a, Barry Starcher b,
Karl Juergen Halbhuber c, David P. Martin d, Ulrich A. Stock a
a
Department of Cardiothoracic and Vascular Surgery, Friedrich-Schiller-University, Bachstr.18, 07743 Jena, Germany
b
Department of Biochemistry, University of Texas, Health Center at Tylers, Tylers, TX, USA
c
Institute of Anatomy II, Friedrich-Schiller-University, Jena, Germany
d
Tepha Inc., Cambridge, MA, USA
Received 29 December 2003; received in revised form 20 April 2004; accepted 5 May 2004
Time for primary review 26 days
Abstract
Objective: To study autologous tissue engineered blood vessels (TEBV) in the descending aorta of juvenile sheep. Methods: Autologous
vascular smooth muscle cells (vSMC) and endothelial cells were obtained from ovine carotid arteries. vSMC were seeded on bioresorbable
scaffolds and dynamically cultured for 14 days. Following endothelialization an additional external ovine small intestinal submucosa
wrapping was applied. Constructs were implanted in the descending aorta of juvenile sheep and removed after 1, 3, 6, 12 and 24 weeks for
evaluation with histological, microscopical and biochemical techniques. Results: Up to 3 months after implantation, grafts were fully patent,
without any signs of dilatation, occlusion or intimal thickening. Scanning electron microscopy revealed a confluent luminal endothelial cell
layer. In contrast, the 6 months graft displayed significant dilatation and partial thrombus formation. Histology displayed layered tissue
formation resembling native aorta. Extracellular matrix (ECM) stains, immunostaining and Transmission Electron Microscopy (TEM)
revealed alternating layers of vSMC and extracellular matrix consisting of collagen, elastin and glycosaminogycans. Compared to native
aorta, the elastin content of the TE grafts was significantly reduced. Conclusion: In this study, we report for the first time, the implantation of
a TEBV in the descending aorta in a large animal model. TEBV were fully functional for up to 3 months. At 6 months the graft remained
functional but significantly dilated, most likely caused by an insufficient elastic fiber synthesis. Hence, future studies need to focus on the
stimulation of elastin synthesis in TEBV.
D 2004 European Society of Cardiology. Published by Elsevier B.V. All rights reserved.
Keywords: Tissue engineering; Aorta; Elastin; Endothelial cells; Vascular smooth muscle cells
1. Introduction
Cardiovascular disease is the leading cause of death in
industrialized countries. For treatment of coronary artery
and peripheral vascular disease, large quantities of blood
vessel substitutes are required. Currently, human autografts (vein or arteries) or prosthetic materials such as
Dacronk are used for this purpose. However, these
* Corresponding author. Tel.: +49-3641-934814; fax: +49-3641933441.
E-mail address: [email protected] (F. Opitz).
substitutes are either not able to grow and remodel or
are prone to material related complications like stenosis,
thrombosis, calcification and infection [1,2]. For this
reason the search for improved blood vessel substitutes
remains unchallenged. Biological alternatives to the currently used prosthetic materials are tissue engineered (TE)
autologous, living blood vessels substitutes. These substitutes have the potential to grow and remodel without
any immunological complications. During the last years
several promising attempts to tissue engineer blood vessels substitutes have been reported. In general they
combine techniques from biological, engineering and
medical sciences. In vitro and in vivo studies have
0008-6363/$ - see front matter D 2004 European Society of Cardiology. Published by Elsevier B.V. All rights reserved.
doi:10.1016/j.cardiores.2004.05.002
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F. Opitz et al. / Cardiovascular Research 63 (2004) 719–730
proven the general feasibility of this approach [3– 7].
However, several biological limitations need to be overcome before tissue engineered blood vessels substitutes
(TEBV) become an alternative to the currently used
prostheses.
The ultimate goal of tissue engineering is to produce
blood vessel substitutes identical to their native counterpart. Native blood vessels consist of several layers of
vascular smooth muscle cells embedded into an extracellular matrix and a luminal endothelial cell layer. The
endothelial cell layer prevents thrombus formation and
the media is responsible for contraction and mechanical
stability of the blood vessel. Probably the most crucial
issue in previous studies was the long-term patency of
the constructs. TE blood vessel substitutes often displayed pathologically increased extracellular matrix
(ECM) deposition and intimal thickening [7,8]. A phenomenon probably caused by excessive proliferation and
secretional activity of vascular cells. Recently we described the in vitro construction of ovine aortic blood
vessels substitutes using autologous, enzymatically derived vascular smooth muscle and endothelial cells [9].
In the present study these blood vessels substitutes were
evaluated after up to 6 month in vivo using histology,
biochemical and ultrastructural methods to determine their
functionality and the extent to which they resembled
native aorta.
2. Materials and methods
2.1. In vitro culture of blood vessel substitutes
Tissue engineered aortic blood vessels substitutes were
created as described previously [9]. In brief, sections of
carotid arteries were harvested from 6- to 8-week-old
lambs. vSMC and EC were isolated by enzymatic dispersion. vSMC were characterized by expression of alphasmooth-muscle-actin (Sigma, München, Germany), basiccalponin (DAKO Cytomation, Hamburg, Germany), heavy
weight caldesmon (Sigma), smooth muscle myosin (Sigma) and the absence of von Willebrand factor (DAKO).
EC were characterized by positive immunostaining for
vWf and negative staining for alpha smooth muscle actin
(ASMA). Cells were cultured in Dulbecco’s modified
Eagle’s medium (DMEM) (Invitrogen, Carlsbad, CA,
USA) containing 10% fetal calf serum (Biochrom, Berlin,
Germany) on tissue culture treated plastic ware (BD
Biosciences, San Jose, CA, USA). Cells were passaged
every 4 days. Approximately 23 ( F 2 days) days were
required to obtain 15 confluent T75 tissue culture flasks.
Subsequently, vSMC were sequentially seeded onto a
custom made poly-4-hydroxybutyrate (P-4-HB) (Tepha,
Cambridge, MA, USA) scaffold using a roller mixer
[9,10]. Seeded grafts were subjected to increasing laminar
flow in a pulsatile bioreactor. Two days prior to implan-
tation TE blood vessels substitutes were endothelialized
with EC.
2.2. Preparation of small intestinal submucosa
In order to stimulate angiogenesis and increase mechanical stability, TEBV were wrapped into ovine small intestinal submucosa (SIS) prior to implantation. Segments of
ovine small intestine (approximately 20 cm in length) were
obtained from a local farm. Small intestines were cleaned by
throughly washing with PBS (Invitrogen). This was followed by incubation in povidone– iodine-solution (Mundipharma, Limburg, Germany) solution for 15 min.
Decontaminated segments were repeatedly washed with
PBS and the mucosa and any connective tissue was removed mechanically [11]. The remaining SIS was transferred into PBS containing an antibiotic solution (1.2 mg
amikacin, 3 mg flucytosin, 1.2 mg vancomycin, 0.3 mg
ciprofloxacin, 1.2 mg metronidazol in 1 ml aqua ad inject.)
and stored at 4 jC. Immediately prior to implantation,
seeded TEBV were wrapped with the decellularized SIS.
2.3. Graft implantation
All animals received humane care in compliance with
the ‘‘Guide for the Care and Use of Laboratory Animals’’
published by the National Institutes of Health (NIH
publication no. 85-23, revised 1985). Autologous grafts
were implanted into the descending aorta of juvenile sheep
replacing an approximately 4 cm long segment. No
postoperative anticoagulation was administered. After 1,
3, 6, 12 and 24 weeks in the animals the TEBV were
explanted. An adjacent segment of native aorta was
explanted for comparision. Explanted tissues were processed immediately or frozen in liquid nitrogen and stored
at 80 jC until analysis.
2.4. Histology
Tissue samples were fixed in 4% paraformaldehyde in
cacodylatebuffer for at least 72 h. Subsequently, samples
were embedded in parafin and 5 Am sections were cut and
stained using H&E, Hart’s, Movat Pentachrom, van Kossa
and Resorcin-Fuchsin staining as described previously
[12].
2.5. Protein extraction
Tissue samples were frozen using liquid nitrogen and
homogenized in a mortar. Protein extraction buffer was
added to the sample with subsequent incubation on a vortex
shaker for 2 h at 4 jC. Tissue debris were pelleted by
centrifugation. The supernatant was collected and protein
content was determined using a commercially available
protein kit (Biorad, Hercules, CA). Samples were shockfrozen and stored at 80 jC for subsequent analysis.
F. Opitz et al. / Cardiovascular Research 63 (2004) 719–730
2.6. Desmosine quantification
Desmosine content of samples was determined using a
radioimmunoassay for desmosine as described previously
[13]. From each vessel three samples were analyzed.
2.7. Immunohistology
Paraffin imbedded tissue samples were cleared by washing with Xylene followed by gradual washing with ethanol
(100%, 80%, 50%). Paraffin free sections were incubated in
blocking buffer (PBS containing 0.1% Triton-X-100 and 2%
skim milk) for 1 h. Slides were washed with PBS-T (PBS
containing 0.1% Triton-X-100). Collagen type I, III and IV
antibodies (all DAKO), Elastin, smooth muscle myosin,
alpha smooth muscle actin and desmin (all Sigma) were
diluted according to instructions in PBS and 150 Al antibody
solution was applied to the sections. Slides were incubated
in a moist chamber for 2 h at 37 jC and then washed
repeatedly with PBS-T. FITC conjugated secondary antibodies (diluted in PBS containing DAPI) were applied
followed by incubation at 37 jC for 1 h. Samples were
washed repeatedly with PBS-T and analyzed immediately
using a fluorescence microscope (Axiovert S 100, Zeiss,
Jena, Germany).
2.8. Electron microscopy
Specimens for transmission electron microscopy (TEM)
were fixed at room temperature in 2% glutaraldehyde
(Sigma) in cacodylate-buffer for 30 min. Specimens were
further fixed in osmium tetroxide (Sigma), dehydrated in
graded alcohol and embedded in araldite. Ultrathin sections
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were stained with lead acetate and lead citrate (Serva).
Samples were investigated on a transmission electron microscope (EM 900, Zeiss, Oberkochen, Germany).
Specimens for scanning electron microscopy (SEM)
were fixed in 4% paraformaldehyde for 24 h. Fixed specimens were dehydrated in an alcohol gradient (40%, 60%,
80% and 99.9%, 10 min each). Samples were critical point
dried using a critical point drier CPD030 (BAL-TEC,
Balzers, Liechtenstein) and mounted on stubs for SEM.
Mounted samples were subsequently coated with gold using
a sputter coating device SCD005 (BAL-TEC). Samples
were analyzed using a scanning electron microscope LEO1450 VP (LEO, Oberkochen, Germany).
2.9. Near infrared femtosecond multiphoton laser scanning
microscopy
Near infrared femtosecond multiphoton laser scanning
microscopy (LSM) was performed as previously described
[14]. Immediately after dissection, native and TE specimens
were investigated at wavelengths of 760 and 840 nm to
induce ECM component dependent autofluorescence as
well as second harmonic generation (SHG) [15]. Owing
to the high near infrared penetration depth, non-invasive
serial optical horizontal sections of the different vessel
specimens were taken in z-steps of 5 and 10 Am within a
depth of 0 –100 Am.
2.10. Statistical analysis
An unpaired t-test was used to determine significant
deviations. p-Values smaller than 0.05 were considered to
be significant.
Fig. 1. Macroscopical appearance of TE blood vessel grafts prior to implantation and 1, 3, 6, 12 and 24 weeks after implantation.
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3. Results
In addition increased thrombus formation was observed in
the dilated graft.
3.1. Macroscopical appearance
3.2. Histology
After one week in vivo the graft displayed a shiny
surface with some irregular regions due to protruding
polymer (Fig. 1). Up to 6 weeks after implantation grafts
were completely patent and displayed no signs of dilatation.
The internal surface was smooth and shiny without any
evidence of thrombus formation. Anastomoses were smooth
and not stenosed. On the external side of the blood vessel
distinct in-growth of small blood vessels into the SIS-layer
was visible. After 12 weeks in vivo, vessels were still
completely patent and displayed no signs of intimal thickening, dilatation or stenoses. However, two small areas of
thrombus formation were observed. In contrast, the graft
cultured for 24 weeks was still patent but clearly dilatated.
3.2.1. Luminal surface
SEM and immunohistology confirmed a confluent endothelial cell lining. SEM revealed a confluent layer of
cells with cobblestone morpholgy (Fig. 2A). A monolayer
of vWf positive and ASMA negative cells was detected in
histolgy sections (Fig. 2B). In addition the luminal cell
layer stained positive for the endothelial cell marker CD31
(data not shown). Cells in the TEBV were more rounded
than native endothelial cells. As observed in the native
aorta ECs aligned parallel to the direction direction of
blood flow. TEM of luminal cells revealed features typical
for endothelial cells (Fig. 2C). Cells displayed peripheral
Fig. 2. Scanning electron micrograph of the luminal surface from native aorta (A) and TEBV cultured for 6 weeks in vivo. Immunostaining of sections from the
same samples using antibodies against vWf (B). TEM images of ECs in the luminal region (C).
F. Opitz et al. / Cardiovascular Research 63 (2004) 719–730
chromatin condensation in the nucleus and many membrane bound cytoplasmatic vesicles. In line with the
macrosopic findings the dilated graft displayed massive
deposition of fibrilar structures and thrombus on top of the
endothelial cell layer.
3.3. Histology
H&E staining revealed a major increase in cell layers
during the first 3 weeks and a slower increase during the
following weeks (Fig. 3). Tissue was arranged in layers with
alternating layers of ECM and perpendicular aligned vSMC
resembling the gross appearance of native ovine aorta.
Replacement of polymeric material with tissue was observed during in vivo culture. No signs of inflammation
such as leukocyte infiltration or ‘‘giant cells’’ were observed
in histology sections.
Resorcin-Fuchsin as well as Hart’s staining revealed a
significant lack in elastin structures in TEBV (Fig. 4A,B).
Compared to native aorta only a fraction of elastic fibers
was detected in TE-vessels. The deminished amount of
elastic fibers in the TEBV was observed throughout the
entire in vivo culture period. Most of the elastic structures
were found in the luminal part of the media. In contrast,
native aorta displayed distinct elastic lamellae throughout
the arterial wall. These findings were confirmed by immunostaining with Elastin specific antibodies (Fig. 5D). Native aortas stained strongly positive for elastin. Elastin was
confined to the predominantly accelular ECM layers. In
contrast, elastin staining was much less distinct in TEBV
and its distribution was more diffuse. Staining of cellular
723
regions in the TEBV suggested ongoing (tropo-)elastin
synthesis.
Movat pentachrome staining revealed that collagen and
ground substance (GAG, Fibronectin) were the predominant
matrix components of TEBV (Fig. 4C). Native aorta also
displayed the presence of collagen and GAG but elastic
fibers appeared to be present in the highest concentration.
Immunostaining revealed that collagen consisted of Collagen type I and Collagen type III fibrils as well as more
difusely distributed Collagen IV in the TE-vessels and in
native aorta (Fig. 5A –C). In contrast, immunostaining for
fibronectin was much stronger in the TEBV than in native
aorta (Fig. 5F).
3.4. Desmosine quantification
Desmosine, which is a an amino acid cross-link specific
to elastin, was quantitated as a measure of elastin content.
The desmosine content in TEBV was significantly reduced
compared to native aorta (Fig. 6). Although the desmosines
increased significantly from week 1 (81.0 pmol/mg protein)
to week 3 (193.7 pmol/mg protein) after three weeks no
further increase could be detected and values remained
significantly lower than in native aorta (4271.0 pmol/mg
protein).
3.5. Ultrastructure
TEM of TEBV revealed a layered arrangement of cells
and extracellular matrix (Fig. 7A). Striated collagen fibrils
were the major components of the ECM in TE blood
Fig. 3. H&E stains of native aorta and TEBV 1,3,6, 12 and 24 weeks after implantation. Note polymeric scaffold residues in the outer layers of the constructs.
Most of the P-4-HB was dissolved during the embedding process resulting in empty spaces in these regions.
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Fig. 4. Resorcin-Fuchsin and Harts staining of native aorta compared to TEBV 12 weeks after implantation (A), van Kossa staining of native aorta and TEBV
24 weeks after implantation (B), Movat Pentachrome staining of native aorta compared to TEBV 6 weeks after implantation (C).
vessels. Elastin was observed as well but to a much lesser
extent than collagen fibrils. In contrast the major ECMcomponent of native aorta was elastin. Striated collagen
fibrils were also observed in the native aorta however less
abundant than in TEBV (Fig. 7B). Elastic and collagen
fibers were co-localized in the lamella of native blood
vessels. Cells growing in the polymeric region displayed
many cytoplasmatic vesicles indicating ingestion of P-4HB (Fig. 7C).
3.6. Near infrared femtosecond multiphoton laser scanning
microscopy
Near infrared femtosecond multiphoton laser scanning
microscopy was performed with the TEBV cultured for
24 weeks in vivo and native ovine aorta. With near
infrared multiphoton excitation at 760 nm, a distinct elastic
fiber system was visualized throughout the native specimens (Fig. 8A –C). In contrast, only a few elastin con-
F. Opitz et al. / Cardiovascular Research 63 (2004) 719–730
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Fig. 5. Immunostaining of native aorta and grafts cultured for 12 weeks. Staining using an antibodies against collagen type I (A), collagen type III (B), collagen
type IV, Elastin (D) and Fibronectin (E).
taining structures were visible at the inner surface of the
TEBV (Fig. 8D – F). Excitation at 860 nm revealed that
organized collagenous fibers were dominant in the TEBV.
In contrast, collagenous fibers were less prevalent in native
tissues. These findings correlate well with the (immuno)histologic observations.
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F. Opitz et al. / Cardiovascular Research 63 (2004) 719–730
Fig. 6. Desmosine content (pmol/mg protein) in TEBVsTEBV compared to
native aorta. For each time-point three samples were taken from random
locations of the TEBV. (*) indicate significant ( p < 0.05) deviations. Error
bars indicate standard deviation.
3.7. Expression of vSMC differentiation markers
Both native aorta and TEBV expressed alpha smooth
muscle actin and smooth muscle myosin in the medial cell
layers. While expression of desmin was not observed in the
medial cell layers (Fig. 9). In contrast the luminal endothelial cells did not stain positive for these vSMC marker
proteins (arrows).
4. Discussion
In this study, autologous, living tissue engineered
blood vessels (TEBV) were implanted in the descending
aorta of juvenile sheep and allowed to remain for up to 6
months in vivo. The general feasibility of this approach
and several important issues concerning the in vivo
behavior of tissue engineered blood vessel substitutes
were addressed.
Most previous studies focussed either on implanting
TE blood vessels in the low-pressure circulation or using
very durable scaffold materials [3,7]. However, very
durable scaffolds are problematic since the scaffolds
interfere with tissue formation and can cause inflammatory responses [16]. In contrast to scaffolds like polyhydroxyoktanoates, P-4-HB based scaffolds degrade in vivo
within months [17]. Patch augmentation of the pulmonary
artery with seeded P-4-HB has been performed successfully [18].
In the present study, P-4-HB was used as scaffolding
material to tissue engineer an entire segment of the
descending aorta in a large animal model. In spite of a
rather short in vitro time span prior to implantation (14
days), all blood vessels withstood systemic blood pressure
for up to 3 months. This positive result is most likely
related to the high elasticity of the P-4-HB scaffolds and
the presence of already organized tissue at the time of
implantation. Application of decellularized small intestine
submucosa resulted in additional stability of the construct
and facilitated angiogenesis. As expected, the scaffolding
material was rapidly replaced by tissue and after 6 months
in vivo the scaffold material was almost completely
absorbed. Minimizing the in vitro time span prior to
implantation resulted in the implantation of a rather immature graft. Thus, rapid in vivo maturation was required,
and in fact was observed during the first weeks following
implantation. Histology revealed that after 3 weeks in vivo,
a newly formed tissue with a mulitlayered ECM alternating
with vSMC layers was present. Our results suggested that
tissue formation reached a peak in the first 6 weeks in vivo
and slowed down subsequently. All vessels remained
patent throughout the entire study (24 weeks) displaying
no signs of occlusion or increased wall thickness. Medial
vSMC displayed markers of late vSMC differentiation
indicating a mature differentiated phenotyp, characteristic
for physiological blood vessels. These findings are particularly important since most previous studies observed
massive intimal thickening and pathologically increased
matrix synthesis resulting in early graft occlusion which
was probably caused by dedifferentiation and unlimited
proliferation of medial vSMC [7,8]. The use of homogenous vSMC populations in this study obviously avoids the
problem of intimal thickening.
In contrast to studies with tissue engineered structures
based on xenogenous and acellular matrix, no calcification
was observed. Recent studies linked calcification to the
presence of xenogenous material (cellular debris, xenogenous ECM). The use of a polymeric scaffold obviously
avoids this problem. Unlike other studies in which polymeric materials such as PGA were employed no inflammatory response was observed in the P-4-HB based
scaffolds.
Inspite of these positive results, a limitation concerning
the ECM structures became apparent in the dilated graft at
24 weeks. The ECM of TEBV were arranged in layers
similar to native aorta. However, in contrast to native aorta,
elastic fibers were almost absent in the ECM layers and
layer formation was less distinct.
Elastic fibers are crucial components of blood vessels
and particularly in elastic arteries, providing elasticity and
flexibility. Compared to native tissue significantly less
elastic fibers were formed in tissue engineered constructs
and the fibers that were present were poorly organized. The
TE constructs showed a marked decrease in desmosine
content compared to native aorta. This could reflect less
elastin or a reduced level of the actual cross-links. When
considered together with the histological data, the lower
desmosine content probably indicates less elsatin deposited
in the matrix.
Instead of elastic fibers collagen and GAGs and fibronectin were the major components of the ECM. GAG was
dominating during early in vivo culture and was successively replaced by collagen. As described before for native
arteries [19], the dominant collagens were types I and III.
F. Opitz et al. / Cardiovascular Research 63 (2004) 719–730
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Fig. 7. (A) Transmission electron micrographs from native aorta and TEBV cultured for 6 weeks. Location of elastin fibers (E), Collagen fibrils (C) and vSMC
(Ce) are indicated. (B) Enlargements from the ECM layer displaying dominance of elastin in native aorta compared to dominace of collagen in TEBV. (C)
Micrograph from TEBV cultured for 1 week displaying phagocytic vesicles.
In addition collagen type IV, a component of the basal
lamina, surrounded the vSMC. Collagenous fibers still
provided sufficient stability to avoid graft rupture but they
were not able to provide the required elastic properties.
The fact that elastic fiber formation was reduced significantly in all TEBV strongly suggests that the observed
dilatation was not an incidential event. During the first few
weeks following implantation the lack of mature elastic
components was well compensated for by the elasticity of
the polymeric scaffold. However, after degradation of the
scaffold material the newly formed tissue was too weak to
withstand the aortic blood pressure resulting in a dilatation.
The observed lack in mature elastin structures in TEconstructs was described previously [5,6] and seems to
be a serious limitation for tissue engineering functional
TEBV for the high pressure circulation. Synthesis of the
highly cross-linked elastin network seems to be one of the
most complex process during maturation of blood vessels.
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F. Opitz et al. / Cardiovascular Research 63 (2004) 719–730
Fig. 8.
Fig. 9.
F. Opitz et al. / Cardiovascular Research 63 (2004) 719–730
During normal embryogenesis several, complex physical
and biochemical signals are involved in blood vessel
maturation [20]. Medial vascular smooth muscle cells
undergo a sequential differentiation program during embryogenisis [21]. In this study a bioreactor system was used to
mimic the increasing flow conditions. While we observed
that pulsatile flow increased elastin deposition during in
vitro culture, the total amount of elastin was still dramatically reduced [9]. Obviously, additional biomechanical or
biochemical signals are required for a physiological elastin
synthesis.
Hence future studies will need to focus on strategies to
improve mature, cross-linked elastic fiber formation in
TEBV. It has been reported previously that application of
TGF-ß1 and insulin increases elastin snthesis in vitro [22].
Other substances known to increase elastogenesis are calcitriol [23] and retinoic acid [24,25]. The thermoplastic nature
of the P-4-HB scaffolds used in this study allows incorporation of these molecules. Another option is to transfect
vSMC with the lysyl oxidase gene, known to be required for
cross-linking ECM proteins like collagen or elastin (26).
The availibilty of the near infrared laser scanning microscopy technique to monitor elastic fiber formation can
facilitate further studies.
Thrombus formation is another recognized problem in
vascular grafts which is probably related to the presence of
an intact, functional endothelium. In this study TEBV
possessed a confluent endothelial layer consisting of vWf
positive cells. In spite of complete endothelialization, localized thrombus formation was observed after 12 and 24
weeks. While in the latter case this is most likely a
consequence of disturbed flow conditions due to the dilatation, the reason in the 12 weeks graft remains unclear. A
possible clinical solution to this problem may be low dose
anticoagulation therapy after implantation. However, further
studies with larger sample sizes are necessary to study
thrombogenicity and endothelial functionality.
Due to the systemic pressure in the descending aorta, and
the weak mechanical stability of unseeded P-4-HB, it was
not possible to implant any control groups. In a previous
study it was demonstrated that unseeded PHB patches
implanted in the pulmonary artery did not support organized
tissue formation sufficiently [18]. This result was confirmed
by studies with other polymeric scaffolds implanted without
cell seeding [27]. Migration of cells from the adjacent native
aorta in- and onto the TEBV cannot be excluded completely.
However, the equal distribution of tissue over the entire
length of the construct after a short period in vivo strongly
support the assumption that the tissue was indeed derived
from the seeded cells. This assumption is further supported
729
by the fact that vSMC in the TE-grafts resemble carotid
vSMC, which display a more elongated morphology than
aortic vSMC.
In conclusion, this early in vivo study reports for the
first time the implantation of a TEBV in the descending
aorta in a large animal model. The TEBV resembled native
aorta and remained patent throughout the study. However,
after 6 months significant dilatation of the graft was
observed. This was most likely caused by the inability of
the cells in the TEBV to synthesize or assemble the normal
complement of elastic fibers in the new matrix of the
engineered vessels. Future studies with larger sample
groups will have to focus on possibilities to support elastic
fiber formation and the issue of thrombogenicity. Ultimately we envision transfering the technique to small
diameter vessels which are clinically more relevant then
aorta descendens.
Acknowledgements
The authors would like to thank Ilka Degenkolbe for
expert technical assistance. R. Kaiser and M. Beck for
excellent assistance with the electron microscopy, U.
Moeller and S. Hitschke for excellent assistance with the
histological staining and K. Koenig for permission to use
the Laser scanning microscope. This research was supported by grants from the German Heart Foundation (UAS)
and the Interdicsiplinary Center for Clinical Research
(UAS, FO).
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