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Am J Physiol Heart Circ Physiol 288: H1451-H1460, 2005. First published October 14, 2004; doi:10.1152/ajpheart.00479.2004
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Engineering of fibrin-based functional and implantable small-diameter blood vessels

Daniel D. Swartz,1 James A. Russell,1 and Stelios T. Andreadis2

1Department of Physiology and Biophysics and 2Bioengineering Laboratory, Department of Chemical and Biological Engineering, State University of New York, Buffalo, New York

Submitted 20 May 2004 ; accepted in final form 7 October 2004


   ABSTRACT
TOP
 ABSTRACT
MATERIALS AND METHODS
 RESULTS
 DISCUSSION
 REFERENCES
 
We engineered implantable small-diameter blood vessels based on ovine smooth muscle and endothelial cells embedded in fibrin gels. Cylindrical tissue constructs remodeled the fibrin matrix and exhibited considerable reactivity in response to receptor- and nonreceptor-mediated vasoconstrictors and dilators. Aprotinin, a protease inhibitor of fibrinolysis, was added at varying concentrations and affected the development and functionality of tissue-engineered blood vessels (TEVs) in a concentration-dependent manner. Interestingly, at moderate concentrations, aprotinin increased mechanical strength but decreased vascular reactivity, indicating a possible relationship between matrix degradation/remodeling, vasoreactivity, and mechanical properties. TEVs developed considerable mechanical strength to withstand interpositional implantation in jugular veins of lambs. Implanted TEVs integrated well with the native vessel and demonstrated patency and similar blood flow rates as the native vessels. At 15 wk postimplantation, TEVs exhibited remarkable matrix remodeling with production of collagen and elastin fibers and orientation of smooth muscle cells perpendicular to the direction of blood flow. Implanted vessels gained significant mechanical strength and reactivity that were comparable to those of native veins. Our work demonstrates that fibrin-based TEVs hold significant promise for treatment of vascular disease and as a biological model for studying vascular development and pathophysiology.

matrix degradation/remodeling; vascular disease; vascular reactivity; vascular tissue engineering; smooth muscle; endothelial cells

MANY APPROACHES HAVE BEEN taken to replace diseased or damaged blood vessels. Synthetic conduits of polytetra-flouroethylene (Teflon, ePTFE) or polyethylene terephthalate (Dacron) have been used extensively with a great degree of success in replacement of large-diameter (>6 mm) vessels (9). However, small-diameter synthetic grafts displayed high failure rates due to thrombus and plaque formation. Adsorption of proteins or endothelial cells to the luminal surface of the grafts decreased thrombogenicity but did not restore vasoreactivity and long-term patency (8, 9, 11, 26, 37). Allogeneic grafts demonstrated long-term patency and reactivity, but their clinical use is prevented by high immunogenicity. Autografts, predominantly from saphenous veins or radial arteries, are the most widely used for small-diameter vessel replacement procedures such as coronary artery bypass. Although autologous grafts are currently the gold standard, limited availability, especially for repeat grafting procedures, and the pain and discomfort associated with the donor site necessitate the development of alternative technologies.

Tissue engineering approaches that use natural or synthetic biomaterials as three-dimensional scaffolds for cell growth have been proposed. Natural biomaterials, derived from decellularized tissues, demonstrated successful infiltration of host cells and near physiological level of vasoreactivity after long-term implantation in vivo (17). Others have seeded decellularized matrices with endothelial cells and/or myofibroblasts before implantation to improve patency and vasoreactivity (1, 20). A conventional cell culture approach employed sheets of smooth muscle cells (SMCs) and fibroblasts that were manually wrapped around a mandrel to form multilayer tubes (21, 22). This construct had a well-defined layered organization and displayed functional reactivity and burst pressure comparable to native vessels. However, this methodology is very difficult to automate or scale-up and requires a minimum of 12 wk in culture.

Biodegradable scaffolds such as polyglycolic acid (PGA) in combination with application of pulsatile pressure in a bioreactor rendered tubular constructs that demonstrated contractile properties and sufficient strength for implantation in vivo (24, 25). Indeed, when implanted into miniature swine, tissue-engineered arteries exhibited patency that was documented up to 24 days by digital angiography (25). However, recent studies showed that the products of PGA hydrolysis induced dedifferentiation of SMCs in vivo, limiting the efficacy of PGA as a scaffold for cardiovascular tissue engineering (15). Collagen gels have also been used because collagen is a major component of the extracellular matrix of blood vessels and a natural cell substrate (36). When collagen gels were constrained during compaction, both cells and collagen fibrils attained circumferential alignment similar to native tissue (2). Furthermore, application of cyclic strain improved cell alignment, matrix remodeling, and mechanical strength of these constructs (28–30). Even so, collagen gels attenuated collagen synthesis by fibroblasts and SMCs, thus limiting the mechanical strength of the tissues (6, 33). As a result, transplantation of large-diameter (7 mm), collagen-based tissue-engineered blood vessels (TEVs) required reinforcement with Dacron mesh to provide sufficient mechanical strength (16).

Fibrin may be an alternative to collagen as a scaffold for TEV development. Similar to collagen, fibrin gels can achieve high seeding efficiency and uniform cell distribution. Fibrin can be formed by autologous fibrinogen, and the degradation rate of fibrin gels can be controlled by addition of fibrinolytic inhibitors such as aprotinin or {epsilon}-aminocaproic acid (38). In contrast to collagen, fibrin stimulates synthesis of collagen and elastin and yields TEV constructs with improved mechanical properties (14, 23, 35), suggesting that fibrin may be a more appropriate scaffold for cardiovascular tissue engineering.

In this study, we used ovine smooth muscle and endothelial cells to engineer small-diameter (4 mm) blood vessels, which attained considerable mechanical strength and vasoreactivity after only 2 wk in culture. When implanted into 12-wk-old lambs, fibrin-based TEVs exhibited remarkable remodeling with considerable production of collagen and elastin and significantly increased mechanical strength. Furthermore, implanted TEVs exhibited physiological levels of blood flow and vasoreactivity. Therefore, fibrin-based TEVs hold significant promise for treatment of vascular disease and as a model system to address interesting questions with regard to blood vessel development and pathophysiology.


   MATERIALS AND METHODS
TOP
 ABSTRACT
 MATERIALS AND METHODS
RESULTS
 DISCUSSION
 REFERENCES
 
Cell isolation and culture. Ovine vascular SMCs were isolated from umbilical veins of near-term fetal lambs. After removal of excess connective tissue and adventitia, the vessels were cut longitudinally and placed lumen side up. With the use of a scalpel blade, endothelial cells were scraped from the luminal surface with a single pass, and the removed cells were then vigorously pipetted up and down in 1 ml of PBS and placed directly into a T-25 flask with 4 ml of M199 medium supplemented with 15 mM HEPES, 100 U/ml streptomycin, 100 µg/ml penicillin, 1% L-glutamine, and 20% FBS. For SMC isolation, blood vessels were cut into small pieces and placed into a T-25 flask with 3 ml of culture medium. SMCs that migrated from the explants were incubated in M199 medium supplemented with 10% FBS, 100 U/ml streptomycin, 100 µg/ml penicillin, and 15 mM HEPES (all GIBCO, Gaithersburg, MD). Endothelial and smooth muscle cells were labeled by staining with DiI-Ac-LDL and anti-smooth muscle myosin (Biomedical Technologies, Stoughton, MA), respectively. Labeled cells were visualized by fluorescent microscopy, and the purity of cultures was verified using flow cytometry. Cells were used for study at passage 5 or less. SMCs and endothelial cells were incubated with humidified air in 5% CO2 at 37°C. Cells were passed at near confluence with 0.5% trypsin/EDTA.

Preparation of fibrin-based TEV constructs. SMC containing thrombin was mixed with fibrinogen at 1:1 volume ratio resulting in a final cell concentration of 1.66 x 106 cells/ml. The final concentrations were 2.5 mM CaCl2, 2.5 U/ml thrombin, and 3.5 mg/ml ovine fibrinogen (all Sigma, St. Louis, MO). The solution (1.5 ml/tube) was poured into a mold (3-ml syringe barrel) surrounding a 4.0-mm outer diameter silastic tube and gelled within 5–10 s. After 30-min incubation, the mold was removed and the fibrin tube was placed in a 50-ml conical tubes containing 30 ml of culture medium that was supplemented with 50 µg/ml ascorbic acid. The constructs were incubated at 37°C and 5% CO2 for 48 h before addition of aprotinin at the indicated concentrations.

After 2 wk in culture, the constructs, which were to be implanted, were placed into 15-ml conical tubes containing 6 x 106 endothelial cells in 12 ml of culture medium. The conical tubes were capped and placed on a rotating device for 2 h at 6 rpm to facilitate seeding of endothelial cells on the outer surface of the cylindrical constructs. The constructs were then returned to the incubator for 3 or 10 days, at which time they were implanted. Immediately before implantation, the outer surface of each TEV was pulled inward using forceps to place the endothelial cells in the luminal side of the vessel.

Histology and immunohistochemistry. TEV morphology was assessed with standard hematoxylin and eosin staining of paraffin-embedded tissue sections using standard protocols as we described previously (12, 13). To detect collagen and elastin matrix components, paraffin sections (4 µm) were deparaffinized with reverse ethanol-xylene washes and stained with Mason's trichrome and Verhoff's elastin, respectively. von Willebrand factor was used to identify endothelial cells following manufacturer's recommendations (Sigma).

Measurements of reactivity and mechanical strength. TEVs were removed from the silastic tubing mandrel, cut circumferentially in 2- to 3-mm segments, and placed into an isolated tissue bath. The tissue bath contained standard Krebs-Ringer solution that was kept at 37°C and continuously bubbled with 94% O2-6% CO2 to obtain pH 7.4, PCO2 of 38 mmHg, and PO2 >500 mmHg. The Krebs-Ringer solution consisted of (in mmol/l) 118 NaCl, 4.7 KCl, 2.5 CaCl2, 1.2 KH2PO4, 1.2 MgSO4, 25.5 NaHCO3, and 5.6 glucose. The vessels were equilibrated for 30–60 min before a passive tension of 1.0 g was applied. Over the next 60 min, the constructs were rinsed three times and the tissue tension was readjusted to 1.0 g at a stable stretch length. Isometric contraction was recorded by a force transducer (model Statham UC 2, Gould, Cleveland, OH).

Pharmacological agents were added to the bath to elucidate TEV function. Contractions were elicited by adding KCl (118 mM) for 15 min or until tension was stable, norepinephrine (NE; 3 x 10–6 M), or U46619 [GenBank] (10–7 M). For relaxation, the vessels were first contracted with NE (10–6 M) and then relaxed with the sodium nitroprusside derivative (SNAP; 10–7 to 10–6 M). Relaxation was reported as percent of the maximal NE contraction.

Length-tension measurements were collected by incrementally increasing the applied force and recording the new adjusted tissue length. This procedure was repeated until the tissue broke yielding the break tension (or ultimate tensile strength) and break length of the tissue. The initial tissue length corresponds to the length under a passive tension of 1.0 g. Finally, tissue toughness was obtained by numerically integrating the area under the tension-length curves (7) and expressed in units of (mm x g force/g tissue wt; mm g/g).

Implantation of TEV constructs. All procedures and protocols in this study were approved by the Laboratory Animal Care Committee of the State University of New York at Buffalo. Dorset cross castrate males 10–12 wk of age (~25 kg) were fasted 24 h before surgery. Anesthesia was induced with pentathol sodium (50 mg/animal) and maintained with 1.5–2.0% isoflurane through a 6.0-mm endotracheal tube using a positive pressure ventilator and 100% oxygen. The left external jugular vein was exposed through a longitudinal 8-cm incision. After small collateral vessels were tied, 3,000 U of heparin sulfate were administered and the proximal and distal ends of the implantation site were clamped.

The TEV construct was pulled inside-out placing the endothelium to the luminal side of the graft. The external jugular vein was transected, and a 1.0- to 1.5-cm segment of the TEV was sutured into place using continuous running 8-0 proline cardiovascular double-armed monofilament suture (Ethicon, Johnson and Johnson, Somerville, NJ). The vascular clamp was slowly removed, and flow was resumed through the TEV graft. A radiopaque tie was loosely secured at the caudal end of the TEV to mark the location of the graft. The incision was closed using 2-0 vicryl in layers (facia and skin). The animal was recovered and monitored daily for adverse affects. Angiograms and Doppler ultrasounds were performed at 5 and 15 wk postgrafting, respectively. At the indicated times, animals were euthanized using 10 ml of concentrated sodium barbiturate (Fatal Plus, Vortech Pharmaceuticals, Dearborn, MI). TEV grafts were removed along with intact caudal and cephalic native vessel. Tissue segments were processed for histology and reactivity measurements.

Statistical analysis. Results were expressed as means ± SE. Unpaired Student's t-test was performed using Sigma Stat 4.0 software (Jandel), and statistical significance was defined as P < 0.05.


   RESULTS
TOP
 ABSTRACT
 MATERIALS AND METHODS
 RESULTS
DISCUSSION
 REFERENCES
 
Preparation and morphology of fibrin-based TEVs. To prepare vascular constructs, fibrin gels containing ovine SMCs (1.66 x 106 cells/ml) were polymerized around 4.0-mm silastic cylindrical mandrels. The length and outside diameter of TEVs were 30 and 9.0 mm, respectively. Fibrin gels contained 3.5 mg/ml fibrinogen and 2.5 U/ml thrombin. After 14 days in culture, the tubular constructs compacted in the radial and longitudinal directions to 3–5% of starting volume resulting in tissues with high degree of structural integrity (Fig. 1A). The final length and outside diameter of TEVs were 3–6 and ~5.0 mm, respectively. Hematoxylin and eosin stain showed homogeneous cell distribution (Fig. 1B), and trichrome stain demonstrated that the cells synthesized considerable amounts of collagen after 14 days in culture (Fig. 1C).



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  		Fig. 1. Engineering fibrin-based blood vessels. Smooth muscle cells (SMCs) were embedded in fibrin gels and cultured around a 4-mm cylindrical mandrel in the presence of 20 µg/ml aprotinin. A: at 2 wk, tissue-engineered blood vessels (TEVs) were removed from the mandrel and photographed. The tissues contracted significantly in both the radial and longitudinal directions, yielding cylindrical constructs with wall thickness of ~0.5 mm. B: hematoxylin and eosin (H&E) stain of tissue sections revealed circumferential alignment of cells that were close to the lumen but no cellular alignment in the rest of the tissue (magnification x40). Before implantation, endothelial cells (ECs) were seeded in the luminal side of each construct and allowed to grow for 10 days to form a confluent monolayer (inset). C: Mason's trichrome revealed significant deposition of new collagen throughout the tissue. Luminal surface is at the top of B and C (magnification x40).

 
To delay gel degradation, various concentrations of aprotinin, a protease inhibitor, were added to the culture medium. Addition of aprotinin prevented matrix degradation resulting in significantly increased tissue weight and decreased cell density after 2 wk in culture (Fig. 2, A and B). Increasing the concentration of aprotinin above 10 µg/ml did not affect tissue weight or cell density any further.



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  		Fig. 2. Effects of aprotinin on TEV weight and cell density. SMCs were embedded in fibrin gels and cultured around a 4-mm cylindrical mandrel with the indicated concentrations of aprotinin for 2 wk. A: TEV wet tissue weight. B: TEV cell density. Tissue sections were stained with H&E, and the number of cells was counted in 4 randomly selected fields of view. The area of each field of view was also measured using PhotoSpot software, and cell density was calculated as number of cells per surface area (cells/mm2). *P < 0.05 compared with 0 µg/ml aprotinin (n = 3–6). C and D: TEVs were cultured for 8 wk in the presence of 20 µg/ml aprotinin. TEV wet tissue weight (C) and cell density (D) were determined at the indicated times (n = 3–6; except at 1 and 3 wk, n = 2). Data are presented as means ± SE. *P < 0.05 compared with 1-wk time point.

 
Tissues grown in the presence of 20 µg/ml aprotinin maintained their wet weight up to 8 wk in culture, when the experiment was terminated (Fig. 2C). Cell density increased by 30% on the third week and remained constant thereafter (Fig. 2D). By comparison, the steady-state cell density of TEVs (816 cells/ mm2) was 2.5 times less than that of the aorta of 6-wk-old lambs (2,084 cells/mm2; our data).

Mechanical properties of fibrin-based TEVs. To test the mechanical properties of engineered blood vessels, each tissue was cut into small annular rings that were mounted onto a force transducer in an isolated tissue bath. Tissues were stretched gradually, allowing enough time for the force to stabilize before measuring tissue length. This stretching procedure was repeated in small increments until the tissues broke, yielding the values of break tension and break length.

Aprotinin did not affect break tension (Fig. 3A) but increased break length (Fig. 3B) and tissue toughness (integral of tension-length curves; Fig. 3C). At 20 µg/ml aprotinin, tissue constructs attained a significant break tension of 1,700 g/g (g force per g tissue), which was 7% of the break tension of a native carotid artery of 9-mo-old lamb (24,300 g/g; our data). Break length increased from 9.2 to 14.0 mm and tissue toughness increased from 3,866 to 14,941 mm g/g when aprotinin concentration increased from 0 to 20 µg/ml. Break tension and break length did not change significantly over a period of 8 wk in culture, suggesting that fibrin-based TEVs attained considerable mechanical strength early on, obviating the need for long culture times (Fig. 3D).



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  		Fig. 3. TEV mechanical properties. SMCs were embedded in fibrin gels and cultured around a 4-mm cylindrical mandrel with the indicated concentrations of aprotinin for 2 wk. A: break tension at the indicated concentrations of aprotinin. The dashed line represents the break tension of fibrin gels with no SMCs (n = 3–6). B: break length at the indicated concentrations of aprotinin. C: TEV toughness at the indicated concentrations of aprotinin was calculated by numerically integrating the area under the tension-length curves. *P < 0.05 compared with 0 µg/ml aprotinin (n = 3–6). D: break tension and break length of TEV constructs as a function of time in culture. TEVs were cultured for 8 wk in the presence of 20 µg/ml aprotinin, and the break tension and break length of the constructs were determined at the indicated time points (n = 3–6; except at 1 and 3 wk, n = 2). Data are presented as means ± SE.

 
Reactivity of fibrin-based TEVs. The ability of the TEVs to contract or relax was measured after exposure of annular segments of each construct to various vasoactive substances. When exposed to KCl (118 mM), a nonreceptor-mediated vasoconstrictor, TEVs contracted with a maximal force of 60 g/g (no aprotinin), which was significantly higher than the contractile force of a native jugular vein (5.9 g/g; Fig. 4A). Receptor-mediated contraction was also examined using a thromboxane A2 mimetic U-46619 (3 x 10–7 M) and NE (3 x 10–6 M). TEVs demonstrated significant contraction in response to both U-46619 and NE, albeit to a lesser extent than KCl, exerting maximal contractions of 21 and 9 g/g, respectively (no aprotinin). Surprisingly, aprotinin significantly decreased the contractility of TEVs in a concentration-dependent manner for all three vasoconstrictors (Fig. 4A).



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  		Fig. 4. TEV reactivity. SMCs were embedded in fibrin gels and cultured around a 4-mm cylindrical mandrel with the indicated concentrations of aprotinin for 2 wk. A: reactivity in response to KCl (118 mM), U46619 [GenBank] (10–7 M), or NE (3 x 10–6 M). The dashed line represents constriction of native jugular vein in response to KCl (118 mM; n = 3–6). B: TEVs were cultured for 8 wk in the presence of 20 µg/ml aprotinin and reactivity in response to KCl (118 mM), U46619 [GenBank] (10–7 M), or NE (3 x 10–6 M) was determined at the indicated time points (n = 3–6; except at 1 wk, n = 2). C: relaxation in response to sodium nitroprusside (SNAP; 10–7 and 10–6 M). After 2 wk in culture, TEVs were constricted with NE (10–6 M) before addition of SNAP and relaxation was recorded as a percent of NE constriction. *P < 0.05 compared with 0 µg/ml aprotinin (n = 3–6). Data are presented as means ± SE.

 
TEVs remained reactive to KCl for 8 wk in culture. Surprisingly, the change in TEV contractility over time was different for U46619 [GenBank] and NE. During the first 2 wk, TEV contraction had a tendency to decrease in response to NE but increased in response to U46619 [GenBank] , remaining constant for both vasoconstrictors thereafter (Fig. 4B).

To measure relaxation, TEVs were contracted with NE (10–6 M) and relaxation was induced by the nitric oxide (NO) donor SNAP once a plateau level of contraction was achieved. At low SNAP concentration (10–7 M), aprotinin-treated tissues relaxed by ~50% of the NE-induced contraction. Increasing SNAP concentration to 10–6 M induced complete relaxation at all aprotinin concentrations (Fig. 4C). At 200 µg/ml, the percent relaxation was very large because constriction of these tissues to NE was very small (Fig. 4A). However, the absolute value of relaxation to SNAP was not significantly different than relaxation at other aprotinin concentrations (data not shown).

Transplantation of fibrin-based TEVs. For implantation, we made longer TEVs with initial length and outside diameter of 6.0 cm and 9.0 mm, respectively. After 2 wk in culture in the presence of 20 µg/ml aprotinin, the tissues compacted to a length of ~1.5 cm and an outside diameter of ~5.0 mm (Fig. 1A). At this time, endothelial cells were added and the tissues remained in culture for an additional 3 or 10 days before grafting. TEVs containing SMC and endothelial cells were implanted as interpositional grafts into the external jugular veins of 12-wk-old lambs. These allogeneic TEVs were implanted without immunosuppressants but with administration of 60 U/kg heparin.

Angiography at 5 wk postgrafting showed that the grafts (n = 4) were patent with no signs of aneurysm formation (Fig. 5A). At 15 wk postgrafting, the animals (n = 2) were anesthetized and an ultrasound probe was placed around the graft or the control jugular vein on the opposite side of the neck to measure blood flow (Fig. 5, B and C). Blood flow through the graft was 120 ml/min (range between 75–165 ml/min; n = 2 animals), which was comparable (71%) to that of the control side 170 ml/min (range between 120–220 ml/min; n = 2 animals), indicating that the grafts were patent. At that time, the animals were killed and tissues were excised and cut into segments for further examination.



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  		Fig. 5. Implanted TEVs remain patent. TEVs were implanted into the external jugular vein of 12-wk-old lambs. A: representative angiogram (n = 2) at 5 wk postgrafting demonstrates blood flow through the graft. The caudal end of the graft was marked with a radiopaque tie. Contrast was injected from the caudal end of the graft, which diffused retrograde through the graft before clearing by antegrade flow. Ultrasound imaging of the graft (B) and control jugular vein at the opposite side of the neck (C) at 15 wk postgrafting (n = 2). The graft appears to have similar diameter and patency as the control vein. Carotid artery was imaged as a reference to maintain the same orientation of ultrasound images.

 
First, a segment of each TEV implant was cut open for visual inspection. We found that the implants (n = 2) that were cultured for only 3 days after endothelial cell seeding displayed thrombogenicity with minimal space for blood flow. In contrast, the implants (n = 2) that were cultured for 10 days after endothelial cell seeding showed minimal fibrosis and no thrombus formation.

Hematoxylin eosin stain showed that the implanted TEVs integrated well with the native vessels. Similar to native jugular vein (Fig. 6A), explanted TEVs exhibited multiple layers of SMC, which were overlaid by a continuous monolayer of endothelial cells (Fig. 6B and inset). Notably, SMC of implanted TEVs assumed circumferential orientation and produced a high amount of collagen matrix, as revealed by Mason's trichrome staining (Fig. 6, C and D). In addition, implanted tissues exhibited significant amount of elastin, which was organized in fibers (Fig. 6F). However, the fibrillar organization of elastin was not as extensive as that of control jugular vein (Fig. 6E), suggesting that remodeling of implanted TEVs was not complete by 15 wk.



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  		Fig. 6. Remodeling of TEV after implantation. TEVs were implanted into the external jugular veins of 12-wk-old lambs. After 15 wk, the tissues were explanted and processed for histology. A and B: H&E of native jugular vein (A) and TEV graft (B). Inset: staining of ECs with von Willebrand factor (vWF; brown) to demonstrate confluent endothelium at the luminal surface. C and D: Mason's trichrome of native jugular vein (C) and TEV graft (D). Collagen (blue/green) is abundant in both the native vein and the graft. E and F: Verhoff's elastin of native jugular vein (E) and TEV graft (F). Elastin fibers are stained black (arrowheads). In TEVs, elastin appears to be diffuse throughout the SMC layers close to the lumen (#) and elastin fibers are seen at more distant SMC layers (arrowheads). Luminal surface is at the top of A-F, and blood flows in the direction that crosses the plane of the page. Marker represents 100 µm (magnification x20). Cap, capillary.

 
Explanted TEVs exhibited significant contraction in response to KCl. The maximal force of the graft was 6.7 g/g, which was comparable to the maximal force exhibited by the adjacent downstream host vessel (10.8 g/g) and control jugular vein from the opposite side of the neck (5.9 g/g). Finally, the break tension of explanted TEVs increased significantly from 1,700 g/g before implantation to 5,941 g/g, which was 25% of the break tension of native carotid artery of 9-mo-old lambs (24,300 g/g).


   DISCUSSION
TOP
 ABSTRACT
 MATERIALS AND METHODS
 RESULTS
 DISCUSSION
REFERENCES
 
We demonstrated that fibrin may be an appropriate scaffold for SMCs allowing engineering of blood vessels with considerable mechanical strength and vasoreactivity. After a short time in culture, SMCs remodeled the extracellular matrix by substituting the fibrin gel with collagen. Engineered vessels displayed vasoreactivity through receptor- and nonreceptor-mediated pathways and developed considerable mechanical strength that was sufficient for successful implantation. After implantation, TEVs integrated very well with the cephalic and caudal ends of the jugular vein and remodeled successfully producing new collagen and elastin. Explanted TEVs displayed mechanical strength and vasoreactivity similar in magnitude to those of the native vessel. Most importantly, TEVs remained patent and demonstrated blood flow comparable to that of the control (nonoperated) jugular vein, suggesting that fibrin-based blood vessels may provide a promising therapeutic modality and a good model system to study vascular development.

Reactivity of TEV fibrin constructs to various stimuli demonstrated the presence of multiple vasoactive pathways. Vasoconstriction was first measured through a nonreceptor-mediated pathway using KCl, which typically produces maximal contraction. TEVs demonstrated significant reactivity in response to KCl, suggesting functional development. Receptor-mediated contraction through extracellular or intracellular calcium stores was tested using NE and U46619 [GenBank] , respectively. NE acts predominantly through the {alpha}1 voltage-dependent channel receptor that controls the influx of extracellular calcium, whereas U46619 [GenBank] acts on the thromboxane A2 receptor, which is linked to the phospholipase/inositol trisphosphate pathway that controls the release of internal calcium from the sarcoplasmic reticulum (34). TEVs contracted in response to NE and U46619 [GenBank] , indicating significant development of both {alpha}1 and A2 receptor-mediated pathways.

Relaxation of the constructs was measured by using SNAP, a sodium nitroprusside derivative that acts as a NO donor. NO induces relaxation through the soluble guanylyl cyclase (s-GC) pathway, which ultimately activates protein kinase G (27). The TEV constructs were first contracted with NE, and the SNAP response was measured as a percentage of the maximal NE-induced contraction. At a concentration of 10–7 M, SNAP caused a partial relaxation, whereas 10–6 M SNAP caused complete relaxation, indicating that the s-GC pathway was highly functional in TEV constructs.

Our results show that aprotinin had significant effects on multiple aspects of TEV development. First, aprotinin led to a significant decrease in cell density of TEVs. Because aprotinin inhibits fibrinolysis, the decreased cell density may reflect preservation of fibrin matrix rather than decreased cell proliferation. In addition, aprotinin increased break length, toughness, and the ability of TEVs to dilate in response to SNAP. Surprisingly, aprotinin decreased the ability of TEVs to constrict in response to receptor- or nonreceptor-mediated pathways in a concentration-dependent manner, suggesting that tissue reactivity may be independent of mechanical properties. Although the mechanism remains unknown, it may be that inhibition of fibrinolysis prevents development of tissues with high matrix density, ultimately affecting the type and magnitude of mechanical forces that TEVs generate in response to chemical signals. Alternatively, aprotinin may act directly on SMCs by blocking some of the pathway(s) that are necessary for influx of calcium and contractibility. Our current data do not distinguish between these two hypotheses.

Interestingly, TEVs demonstrated significant mechanical strength and physiological levels of vasoreactivity without application of mechanical force, which was previously shown to induce cellular alignment and improve the mechanical properties of collagen and PGA-based TEVs (18, 25, 28–30). Such improvement was at least partially due to a delicate balance between synthesis and degradation of extracellular matrix stimulated by application of cyclic strain (29, 30). More recently, it was demonstrated that cyclic mechanical strain could increase the toughness of TEVs made with a mixture of collagen and fibrin matrices (7). Based on these studies, we expect that application of mechanical forces during development of fibrin-based TEVs may further increase their reactivity and mechanical strength by regulating matrix remodeling.

Fibrin-based TEVs compare favorably with previously reported vascular constructs. Collagen possessed inadequate mechanical properties to withstand implantation. Use of a support sleeve from cross-linked collagen increased mechanical strength, but chemical cross-linking with glutaraldehyde decreased viability of cells entering the sleeve (3). On the other hand, PGA-based TEVs were cultured for 12 wk before they developed adequate strength for transplantation (24, 25). Even after 8 wk in culture, vascular reactivity in response to 10–6 M endothelin was 0.3 g, compared with fibrin-based TEVs, which demonstrated 1.3 g of reactivity in response to KCl after only 2 wk in culture. Fibrin-based TEVs also demonstrated patency and significant remodeling in vivo as evidenced by production of collagen and elastin fibers. The time and extent of remodeling were similar to those observed with vascular constructs made from biodegradable polymeric materials (25, 31). Therefore, fibrin may be a superior matrix for engineering vasoreactive and transplantable vascular constructs.

The fast development of fibrin-based TEVs may be attributed largely to the characteristics of the fibrin matrix. Vascular SMCs bind to fibrinogen and fibrin through integrin {alpha}v{beta}3, eliciting a series of biological responses such as secretion of proteases that degrade the fibrin matrix and production of extracellular matrix, e.g., collagen (5, 19). Recent findings show that fibrin promotes collagen and elastin production to a larger extent than collagen gels (14, 23, 35). Ultimately, the interaction of SMCs with the fibrin matrix results in matrix remodeling and development of mechanical and contractile properties. Fibrin also promotes angiogenesis through attachment and proliferation of endothelial cells (10) but does not support platelet adhesion (4, 32), suggesting that fibrin may be less thrombogenic than other biomaterial substrates.

Indeed, our findings show that implanted TEVs maintained a uniform endothelial monolayer and remained patent for 15 wk in vivo. Angiograms at 5 wk and ultrasound and blood flow measurements at 15 wk postgrafting demonstrated graft patency and blood flow rates comparable to those of the control jugular vein. Notably, explanted TEVs displayed reactivity values that were similar to those of control jugular vein. Finally, histology showed a seamless integration of the graft with the adjacent natural vessel and significant production of collagen and elastin. Collectively, all these data suggest that fibrin-based TEVs are sufficiently strong for in vivo implantation and that they are capable of in vivo remodeling with no detectable immunogenic response or production of toxic by-products.


   FOOTNOTES
 

Address for reprint requests and other correspondence: S. T. Andreadis, Bioengineering Laboratory, 908 Furnas Hall, Dept. of Chemical and Biological Engineering, State Univ. of New York, Buffalo, NY 14260 (E-mail: sandread{at}eng.buffalo.edu)

The costs of publication of this article were defrayed in part by the payment of page charges. The article must therefore be hereby marked "advertisement" in accordance with 18 U.S.C. Section 1734 solely to indicate this fact.


   REFERENCES
TOP
 ABSTRACT
 MATERIALS AND METHODS
 RESULTS
 DISCUSSION
 REFERENCES
 

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