BIOMATERIALS GROUP

TISSUE ENGINEERING OF BLOOD VESSELS

Deniz YÜCEL

Blood vessels function to carry blood from the heart to the tissues and organs, and from the tissues and organs to the heart. They form a branched system of arteries and veins that vary in size, mechanical properties, biochemical and cellular content, and ultrastructural organization, depending on their location and specific function. The largest arteries, such as the aorta, function to transport blood originating from the heart. The smaller diameter muscular arteries deliver the blood from the large arteries to the tissues and organs. These then branch into smaller arterioles and capillaries, which function to distribute blood within the tissue and organs. Blood is returned to the heart through venules, which combine to form veins. (Fig.1)(10)

The large and medium sized arteries have distinct structural features, primarily the intima, media and adventitia, although these are less obvious in the smaller arterioles and do not exist in the capillaries. The intima forms the layer closest to the blood flow and consists of a lining of endothelial cells attached to a connective tissue bed of basement membrane and matrix molecules. This is adjacent to the internal elastic lamina, which is a band of elastic tissue, found most prominently in the larger arteries, and which separates the intima from the media. The media contains a dense population of smooth muscle cells, organized concentrically, with bands or fibers of elastic tissue. The external elastic lamina separates the media from the adventitia, which contains a collagenous extracellular matrix containing fibroblasts, blood vessels and nerves, and functions to add rigidity and form to the blood vessel. The veins are thin-walled vessels, lack the distinct molecular and tissue organization of arteries, and deform more easily. (Fig.1)(10)

The extracellular matrix surrounding the vascular cells is complex and combines to provide the biomechanical properties of the tissue. The molecular network consists primarily of collagen (primarily types I and III), elastin in the form of fibers, proteoglycans (including versican, decorin and biglycan, lumican and perlican), hyaluronan, glycoproteins (for example, laminin, fibronectin, thrombospondin and tenascin). The mechanical properties critical to blood vessel function include the tensile stiffness, elasticity, compressibility and viscoelasticity. The collagens provide the tensile stiffness, the elastin provides the elastic properties, proteoglycans contribute to the compressibility, and combined with the collagen and elastin they are responsible for the viscoelastic properties. This complex mixture of molecules and their organization provide the blood vessels with their properties that allow them to function throughout life. (10)

Figure 1. Human circulatory system, showing some of blood vessels and their structures. (Human Anatomy Fifth Edition Van De Graff )

The improper functioning of blood vessel results from the atherosclerotic lesion consists of a raised focal plaque within the intima consisting of a lipid core, surrounded by extracellular matrix and smooth muscle cells and covered by fibrous cap. As it increases in size through intimal hyperplasia it restricts blood flow and eventually blocks vessels. Atherosclerotic vascular disease, including peripheral vascular and coronary artery disease, is the major cause of mortality and morbidity in the United States, Europe, and other western nations. Current surgical therapy for diseased vessels less than 6 mm in diameter involves bypass grafting with autologous arteries or veins, and coronary artery bypass grafting procedures are performed approximately 600,000 times annually in the United States alone. Although common surgical practice, vascular grafting has significant limitations and complications. Arterial conduits have restricted dimensions and are limited in supply. Venous conduits lack vasomotor tone and may have varicose degenerative alterations that can lead to aneurysm formation in the higher-pressure arterial circulation. Allografts are problematic because of a high rate of rejection. Synthetic materials are excessively thrombotic when used to bypass arteries less than 6 mm in diameter, with thrombosis rates higher than 40% after 6 months. (1) The cellular and humoral responses to synthetic materials include the deposition of plasma proteins and platelets, the infiltration of neutrophils and monocytes, and the migration of endothelial and smooth muscle cells. Even the most inert substances thus so far developed are still noticed as “foreign”. The tissue/graft/blood interfaces are highly complex microenvironments that are ultimately responsible for graft patency. (3)

The deposition of plasma proteins and platelets: Serum proteins are adsorbed to the graft wall immediately after blood flow is restored. Albumin, fibrinogen, and IgG, the most abundant serum proteins, adsorb to the graft almost instantaneously following exposure to the systemic circulation. Since platelets and blood cells interact predominantly with the bound proteins and not the prosthetic material itself, the concentration of surface ligands presented by the bound proteins has significant short and long-term effects on graft survival. Thus, platelet deposition and activation is influenced by prior protein adsorption. Early platelet deposition occurs primarily via receptor mediated interactions with adsorbed proteins and, to a lesser extent, by direct adherence to the graft. The platelet/ protein complex is mediated through von Wilebrand factor and platelet membrane glycoproteins. After adherence, the platelets undergo conformational changes and degranulate, thereby releasing a variety of bioactive substances, including serotonin, epinephrine, ADP, and thromboxane A2, which activate additional platelets and increase thrombin production. The role that acute platelet deposition has on late graft thrombosis and failure remains unclear (3).

The infiltration of neutrophils and monocytes: The acute inflammatory response is mediated through potent chemoattractants like C5a and leukotriene B4, which draw neutrophils to the synthetic graft surface. Activated neutrophils release oxygen free radicals and various proteases, which result in matrix degradation and may inhibit complete endothelialization and tissue incorporation of the vascular graft. (3)

Endothelial Cell (EC) and Smooth Muscle Cell (SMC) ingrowth: Native uninjured blood vessels possess an endothelial lining which constantly secretes bioactive substances which inhibit thrombosis, promote fibrinolysis, and inhibit SMC proliferation to help maintain normal blood flow. Unlike animal models in which there is often complete endothelialization of synthetic grafts, grafts implanted into humans manifest only limited endothelial cell ingrowth not extending beyond 1-2 cm of both anastomoses. Injured or subconfluent ECs may be found in the perianastomotic region, are phenotypically altered, and may promote thrombogenesis and SMC growth. Subintimal SMC proliferation and intimal thickening occur predominantly at the perianastomotic region, which are areas of chronic endothelial cell turnover as well as areas of complex biomechanical characteristics and chronic inflammation. Smooth muscle cells, inflammatory mononuclear phagocytes, and foreign body giant cells also produce a variety of growth modulating substances including PDGF, FGF, and TGF-β, which may perpetuate SMC proliferation and production of extracellular matrix components. This hyperplastic SMC response is thought to be a major contributor to the development of intimal hyperplasia. (3)

The chemical composition, construction parameters and biomechanics of a vascular graft influence its interaction with its host. The differences of graft materials, porosity, compliance, electrical charge and surface texture all contribute to the magnitude and characteristics of the body’s inevitable “foreign body” reaction. The rate of tissue ingrowth is dependent upon graft porosity; however, high porosity grafts require preclotting. In addition, transinterstitial ingrowth may differ comparing to materials with equal porosity, e.g., PGA vs Dacron. (3)

As a result of these complications, the need for a tissue-engineered vessel of small caliber composed of biological materials and autologous cells has arisen and has been an area of active investigation for more than 15 years. (1)

Tissue Engineering:

Replacement with mechanical devices or artificial organs is limited by an increased risk of infection and thromboembolism and finite durability. Because of the above shortcomings, the field of tissue engineering and selective cell transplantation was born as a means to replace diseased tissue with living tissue that is "designed and constructed to meet the needs of each individual patient". The goal of tissue engineering is to "restore, maintain or improve tissue function through the delivery of living elements which become integrated into the patient". The three major approaches include guided tissue regeneration using engineered matrices alone, the injection of allogenic or xenogenic cells alone, or the use of cells placed on or within matrices. Using isolated cells or cell substitutes avoids potential surgical complications and allows cell manipulation before injection/infusion (such as gene therapy), but has the drawbacks of possible rejection or loss of function. The use of cell-matrix constructs, the most common method in tissue engineering, involves either open or a closed system. An open system begins with the in vitro culture of isolated cells. The cells are then seeded onto a scaffold, either synthetic or natural. After appropriate cultivation time, the cell-matrix construct is implanted into the host. The matrix functions to guide the development of the new tissue and provides structural support. In a closed system, the cells are isolated from the body by a permeable membrane allowing exchange of nutrients and waste but protecting the cells from the immune response. (1)

Polymer scaffolds can be produced from natural or synthetic materials. Natural materials may closely mimic the native cellular environment as they are often extracellular matrix components and include collagen, hydroxyapatite, Matrigel (Collaborative Biomedical, Madison, WI), and alginate, among others. Synthetic materials have the advantage of being able to better control material properties such as strength, degradation time, porosity, and microstructure. Cell attachment can be improved by modifying the polymer chemically or by coating it. Growth factors can also be incorporated into the matrix. Defined shapes and sizes can be fabricated readily and reproducibly. Ideally these polymers must be biocompatible and bioabsorbable, nonimmunogenic, support cell growth, and be able to induce angiogenesis to supply the newly formed tissue. Polyhydroxyalkanoates (PHAs) are a class of natural polymers with thermoplastic properties. They are biocompatible, resorbable, extremely flexible, and induce only a minimal inflammatory response. Because of these properties as well as their tensile strength, they are currently being used in the tissue engineering of heart valves and blood vessels. (1)

Tissue Engineering of Blood Vessels:

Some of the critical issues for tissue-engineered blood vessel substitutes are as follows: (9)

· Immune acceptability, off-the-shelf availability

· Functional three-dimensional construct

· Short in vitro culture period

· Control of biological responses

· Validated test beds to predict clinical efficacy

In the above a critical issue is the engineering of a functional three-dimensional substitute or construct. To do this requires providing the following characteristics: (9)

· Biological scaffold

· Sufficient mechanical strength

· Elastic mechanical properties

· Vasoactive with contractile phenotype smooth muscle cell

· Adherent, quiescent endothelium

Small Vessel Constructs:

Small vessel vascular tissue engineering can loosely be categorized into three groups. These include using nonbiodegradable grafts seeded with cells, using natural materials (such as collagen) as grafts with or without cells, and using biodegradable polymer matrices seeded with cells. (1)

Ø Hybrid vascular grafts, consisting of synthetic material such as Dacron (C.R. Bard, Haverhill, PA) or polytetrafluoroethylene seeded with cultured endothelial cells before implantation, have been investigated by numerous groups. (1) Numerous biological substances have been applied to Dacron and to ePTFE thereby creating “biohybrid” grafts. Various proteins, anticoagulants, and antibiotics have been bonded to synthetic grafts to help prevent graft failure. These surface modifications modulate the body’s acute chronic response to the graft material. Grafts impregnated with albumin were initially created in the early 1970s. Various techniques including alkylation, plasma discharge, and the application of thin polymer films have been utilized in the impregnation process. Knitted Dacron prosthesis coated with albumin, gelatin, and collagen were available for clinical use. The protein coating decreases graft porosity thereby eliminating the need for preclotting. When the impregnated protein is degraded, the graft undergoes tissue ingrowth. The albumin surface may also diminish acute thrombogenicity of the graft by reducing the receptor-mediated platelet binding to fibrinogen. (3)

Ø In 1978, the first successful isolation of endothelial cells from segments of vein and their subsequent transplantation onto synthetic vascular grafts was reported by Herring. Theoretically, the presence of a confluent monolayer of endothelial cells could improve graft thrombo-resistance and prevent the development of pseudointimal hyperplasia. Normal, uninjured endothelial cells possess a negative outer charge that repels platelet adherence. They also express glycosaminoglycans, which bind anti- thrombin III, and produce protacyclins and tissue plasminogen activator, all of which facilitate endothelial cells’ anticoagulant activity. However when perturbed, endothelial cells produce a variety of procoagulants like von Willebrand factor, plasminogen activator inhibitor, thrombospondin and collagen. Endothelial cells are also capable of producing numerous substances that either inhibit or stimulate smooth muscle cells proliferation and therefore may modulate development of pseudointimal hyperplasia and factors inducing either vasorelaxation or vasoconstruction. Endothelial cell seeded grafts have also been shown to be more resistant to bacterial contamination. One of the difficulties with endothelial cell seeding is the relatively low cell density initially applied to the graft and inadequate cell attachment to the graft. In trying to solve these problems researchers have attempted a two stage seeding procedure in which endothelial cells are harvested, allowed to proliferate in vitro, and then seeded and grown to confluence on the vascular graft prior to implantation (3).

Ø Miwa and Matsuda created a graft of polyurethane, an artificial basement membrane composed of a complex gel of type I collagen and dermatan sulfate, and an autogenous endothelial cell monolayer. They implanted these grafts into the carotid arteries of dogs without anticoagulation and noted an overall patency rate of 75% (1).

Ø Weinberg and Bell were the first to construct a tissue-engineered blood vessel composed of natural materials (1).

Method: Bovine aortic endothelial cells, smooth muscle cells, and adventitial fibroblasts were isolated and cultured by standard methods. The middle layer of the blood vessel model, corresponding to the media of an artery, was prepared by casting culture medium, collagen, and smooth muscle cells together in an annular mold. The mixture jelled after a few minutes at 37˚C and contracted within a few days to produce a tubular lattice around the central mandrel. After one week, an open Dacron mesh sleeve was slipped over the lattice to provide additional mechanical support. The outer layer, corresponding to the adventitia, was cast around the first lattice with adventitial fibroblasts rather than smooth muscle cells. Two weeks later, when the outer layer was fully contracted, the tube was carefully slipped off the mandrel with jeweler’s forceps and either used for mechanical testing or lined with endothelial cells. For the lining with ECs, the model was cannulated, a suspension of endothelial cells was injected into the lumen, and the vessel was rotated around the longitudinal axis at 1 rev/min for 1 week to distribute endothelial cells uniformly on the luminal surface (5).

Results: The model grossly resembled a muscular artery, except for the Dacron mesh. Electron microscopy showed that the smooth muscle cells are well-differentiated bipolar cells containing bundles of filaments with dense bodies. They frequently appeared to be secreting collagen into the extracellular space, thus contributing to the matrix. The endothelial cells formed a monolayer of flattened cells with intercellular junctions, numerous vesicles, occasional Weibel-Palade bodies, and patches of basement membrane by 1 week. Scanning electron microscopy showed that virtually the entire luminal surface was covered by endothelial cells (5).

The endothelial lining of the blood vessel model functioned like a normal endothelium in several respects. It produced von Willbrand factor, a widely used marker for vascular endothelium, and it formed a permeability barrier for large molecules such as albumin. Endothelial cells release prostacyclin, which is a potent inhibitor of platelet aggregation and is believed to prevent thrombosis in vivo. (5)

Figure 2. Burst strength of the blood vessel model. A) Burst strength as a function of collagen concentration in the wall. B) Models were tested at various times after the second layer was cast. .Science,vol231. 397-399

The ability of the blood vessel model to withstand intraluminal pressure depended on several factors including the mesh, collagen concentration, initial cell density, and time elapsed after casting. The role of these factors was assessed by measuring the burst strengths of models made by varying these parameters from the standard protocol. A model constructed with three layers of collagen lattice alternating with two meshes had a burst strength of 120 to 180mmHg and usually failed by developing a pinhole leak. The burst strength of the model was proportional to the logarithm of the collagen concentration. (Fig.2) (5)

A blood vessel model attained its maximal burst strength 3 to 6 weeks after casting. The increase in strength with time is probably due to cross-linking of the collagen. The decrease in strength after long times may be caused by collagenase secreted by the smooth muscle cells and fibroblasts in the lattices. (5) These constructs were unable to attain adequate burst strengths for in vivo applications, despite their reinforcement with Dacron mesh (1)

Thus, this model meets many of the physiological and physical criteria for a blood vessel model; however, there are substantial differences between the model and normal arteries in addition to the requirement for the Dacron mesh. A major difference is that researchers are unable to include elastin, the principal arterial connective tissue protein besides collagen, in the matrix mixture, although small amounts of elastin may be synthesized by the smooth muscle cells after long periods in culture. A significant structural difference is that the smooth muscle cells and collagen fibers have a largely longitudinal orientation, because the contraction of the lattice layers around the mandrel is primarily radial rather than in the alternating left-and right-handed spirals of blood vessels. This may explain why a model that was not supported by a mesh failed by splitting lengthwise. Another difference is that the densities of smooth muscle cells and collagen in the model are one-eighth to one-fourth those in normal blood vessels. (5)

Ø L'Heureux and colleagues improved on the mechanical strength of these constructs by alterations in culture conditions (1).

Figure 3. The structure of the tissue engineered blood vessel model. www.fmed.ulaval.ca/loex/FASEB.html

Method: Human vascular smooth muscle cells cultured with ascorbic acid produced a cohesive cellular sheet. This sheet was placed around a tubular support to produce the media of the vessel. At this stage, the construct was placed in a bioreactor designed to provide both luminal flows of culture medium and mechanical support. After a week of maturation, the next step was to roll a sheet of fibroblasts around the vascular media to provide an adventitia. Finally, after a maturation period of at least 8 wk, the inner tubular mandrel was removed and the TEBV was either tested for mechanical strength or cannulated at both ends for luminal endothelial cell seeding. (Fig.3) (2)

Results: Macroscopically, the TEBV appeared as a homogeneous tubular tissue strikingly resembling a human artery (Fig. 4A). Histological analysis revealed well-defined tissues: intima, media, and adventitia (Fig. 4B). The adventitia exhibited very dense collagenous layers as well as abundant fibroblasts. In the media, SMC appeared as elongated cells with circumferential or longitudinal orientations. SMC density, although very high for an in vitro model, was still lower than in normal vascular media. SMC did not penetrate the dense internal membrane even though they were in contact with it for more than 8 wk and exposed to a gradient of nutriments leading to the lumen. Metabolic labeling and immunostaining of the endothelium (Fig. 4C) revealed confluent and active ECs as demonstrated by ac-LDL uptake and by von Willebrand factor, two specific EC functions. SMC stained positively for muscular markers α-smooth muscle actin and desmin (Fig. 4D). Adventitial cells were negative for desmin or α-smooth muscle actin, but fibroblasts expressed vimentin and synthesized high amounts of elastin assembled in small fibers, which were organized in large circular arrays (Fig. 4E). Immunostaining also indicated that the ECM contained type I, III, and IV collagens as well as laminin, fibronectin, and chondroitin sulfate. (2)

Figure 4. Organization of the TEBV. A) Macroscopic view of

a mature TEBV (9 wk of adventitial maturation). The vessel is

self-supporting when removed from culture medium (open

lumen = 3 mm). Note that the various layers now form a

continuous vascular wall (inset, graduation = 1 mm).

B) Paraffin cross section of the vascular wall stained with

Masson's trichrome shows collagen in blue-green and cells

in dark purple. Aside from an oversized internal elastic lamina

(IM=125 µm), the histology is similar to that of a muscular

artery with a large media (M=320 µm) and a surrounding

adventitia (A=235 µm).

C) En face view of the endothelium seeded on the IM.

Cytoplasmic green fluorescence reveals cell viability, metabolic

activity, and degree of confluence. Red fluorescence

(DiI-ac-LDL uptake) confirms cell viability and the endothelial

nature of the cells. Blue fluorescence shows the characteristic

von Willebrand factor expression in ECs (orange = red + green;

pink = red + green + blue). Scale bar = 25 µm.

D) Frozen cross section of the media–adventitia junction (arrows)

stained for desmin (nuclei are stained blue). Scale bar = 50 µm.

E)Frozen cross section of the adventitia double stained

for elastin (green) and vimentin (red). Scale bar = 50 µm.

www.fmed.ulaval.ca/loex/FASEB.html

During the maturation period, the burst strength of the adventitia steadily increased from the 1st to the 7th week, when it reached a plateau at 2232 ± 251 mmHg (Fig.5). This plateau was maintained until at least the 12th week, and some adventitias were matured for 24 week without significant change in strength (2382±249 mmHg). (2)

Figure 5. Mechanical strength of the TEBV.

Burst strength of tissue-engineered adventitia

as a function of maturation time.

www.fmed.ulaval.ca/loex/fig3.html

Consequently, the complete vessel had a burst strength over 2000 mmHg. This is the first completely biological TEBV to display a burst strength comparable to that of human vessels. These structures were implanted into the femoral arteries of mongrel dogs and remained patent in 50% at 1 week. Patent grafts showed a smooth thrombus-free luminal surface and did not show signs of degradation, tearing, or dilatation. In all grafts, intramural blood infiltrations were observed between tissue layers, but these did not correlate with decreased patency; some patent grafts had significant infiltrations (Fig. 6B) whereas some thrombosed grafts had minimal infiltrations. Aside from blood infiltration, histological analysis revealed that the graft's architecture was retained. (2)

Figure 6. In vivo grafting of the TEBV.

A) Angiogram of the lower limbs 7 days after implantation. Two patent TEBV are visible (arrows) providing normal blood flow in both legs B) After angiographic examination, the TEBV was explanted along with adjacent segments of the femoral artery, gently flushed with saline, fixed in formaldehyde, longitudinally slit open, and pinned down to expose the luminal surface. Blood infiltrations are seen inside the vessel wall. The luminal surface is free of thrombus. Anastomosis showed no signs of deterioration. Graduation = 1 mm. www.fmed.ulaval.ca/loex/fig5.html

These results suggest that this novel technique can produce completely biological vessels fulfilling the fundamental requirements for grafting: high burst strength, positive surgical handling, and a functional endothelium. The most critical factor: the time needed for TEBV production. It is obvious that this type of engineered autologous tissue is not designed to be used for emergency surgeries. However, in a nonurgent setting, a significant period of time would be available for the production of a TEBV, especially if this vessel provides excellent long-term patency. Nonetheless, the culture period needed for graft production should be as short as possible for maximal clinical applicability. (2)

Ø Campbell and associates tested the hypothesis that cells from nontraditional sources can be used to create new grafts and achieve secondary function once in the arterial environment. They placed silastic tubing into the peritoneal cavity of rats and rabbits. After 2 weeks, the tubing was covered by multiple layers of myofibroblasts, collagen matrix, and a single layer of mesothelial cells. The tissue was removed from the tubing and everted to place the mesothelial layer within the lumen. The tube of tissue was then grafted into the carotid artery or abdominal aorta of the same animal and remained patent for at least 4 months. This novel approach produced grafts that also responded to contractile agonists in a similar fashion to native vessels (1).

Ø Teebken and colleagues created decellularized matrix tubes by enzymatic cell extraction of porcine aortas and then seeded these natural grafts with human endothelial cells and myofibroblasts. (1) The grafts exposed to pulsative flow conditions to evaluate cell retention and proliferation. A bioreactor is used to mimic in vivo conditions in order to maintain virtually complete endothelialisation before potential implantation. The shear stress provided by the perfusion technique allows cells to change their morphology from round to flat and get their normal in vivo appearance. Samples for histology and scanning electron microscopy were taken after 120min and 4 days of perfusion (8).

After 2 h of perfusion using culture medium, the cells were still attached to the luminal surface of the aortic wall; both round and polygonal flat cells were found. Following 4 days of perfusion in the bioreactor, the endothelial cells had spread and divided on the graft surface. This led to an almost complete coverage of the aortic luminal surface, with an endothelial monolayer and a partial endothelialisation. (8)

This model of a tissue-engineered vascular graft has many characteristics of a normal blood vessel. The absence of synthetic material may allow complete graft integration and provide a truly stable artificial blood vessel. (8)

However, myofibroblasts do not penetrate the matrix after having been seeded on the inner surface. Furthermore, the acellular matrix graft will not elicit a substantial inflammatory reaction that could either damage the graft wall (especially the elastic lamellae), setting the scene for long-term aneurysm formation, or trigger acute thrombosis. (8)

Ø Huynh and colleagues constructed a 4-mm diameter graft from small intestinal submucosa and type I bovine collagen. Small intestinal submucosa is a biomaterial, composed primarily of type I collagen, that has shown good patency as a large diameter graft in the canine aorta. Their results revealed excellent patency up to 3 months; histologically, the grafts were remodeled into cellularized vessels that responded appropriately to vasoactive agents such as norepinephrine, serotonin, and bradykinin (1).

Ø Niklason and colleagues pioneered the approach of combining cells with biodegradable polymer. In this model it was reported that the development of techniques to produce small-caliber autologous arteries in vitro from vascular cells grown on a biodegradable polymer matrix, by means of a pulsatile perfusion system for vessel culture (4).

Figure 7. A perfusion system providing pulsative flow to four bioreactors, each contains one tissue engineered vessel. Gas exchange to perfusate occurs via the reservoir bag and perfusate pressure is monitored continuously. Transplant Immunology 1997;vol5 303-306 (6)

The biomimetric system used for vessel culture is composed of bioreactors containing engineered vessels assembled in a parallel flow system. (Fig.7) A suspension of cultured SMCs isolated from the medial layer of bovine aorta was pipeted onto tubular biodegradable polyglycolic acid (PGA) scaffolds that were secured in bioreactors. The surface of the PGA scaffolds was chemically modified with sodium hydroxide, which caused ester hydrolysis on the surface of the fibers, leading to increased hydrophilicity, increased adsorption of serum proteins and improved SMC attachment. After an initial SMC seeding period of 30min, the bioreactors were filled with medium and the SMCs were cultured under conditions of pulsatile radial stress for 8 week. Control vessels were cultured without pulsatile radial stress under otherwise identical conditions. (4)

Results: After 8 weeks of culture, the gross appearance of the vessels was identical to that of native arteries. Histologic examinations of pulsed vessels revealed that SMCs migrated inward to envelop PGA fragments in the vessel wall, resulting in a smooth luminal surface onto which bovine aortic endothelial cells could easily be seeded (Fig.8A and B). In contrast, nonpulsed controls exhibited no such inward SMC migration through the polymer scaffold (Fig.8C and D) and possessed an uneven layer of polymer fragments in the vessel lumen. Thus, vessels cultured under pulsatile conditions had a histologic appearance more similar to that of native arteries. (4)

On the basis of these observations of vessel lumen morphology, bovine aortic EC layers were applied to vessels that had been cultured for 8 weeks under pulsatile conditions. After EC seeding, continuous perfusion of the vessel lumens was instituted for the final 3 days of culture. After 3 days, the presence of an endothelial layer on the luminal surface was confirmed by scanning electron microscopy (Fig.8G) and by staining for von Wilebrand factor and platelet endothelial cell adhesion molecule (PECAM Fig.7H). Pulsed constructs had rupture strengths of more than 2,000 mm Hg, adequate suture retention strength, and collagen content of 50%. These tissue-engineered arteries were implanted into Yucatan pigs with patency demonstrated up to 4 weeks. (4)

Figure8. Photomicrographs of engineered vessels. (A) Pulsed vessel cultured for 8 weeks. (Verhoff's stain for elastin; ×20 before 33% reduction.) (B) Pulsed vessel cultured for 8 weeks. #Cellular region, *polymer region. (Masson's trichrome stain for collagen; ×100 before 33% reduction.) (C) Nonpulsed vessel cultured for 8 weeks. (Verhoff's stain; ×20 before 33% reduction.) (D) Nonpulsed vessel cultured for 8 weeks. (Masson's trichrome stain; ×100 before 33% reduction.) (E) Pulsed vessel without medium supplementation. (Verhoff's stain; ×20 before 33% reduction.) (F) Pulsed vessel without medium supplementation. (Masson's trichrome stain; ×100 before 33% reduction.) (G) Scanning electron microscopy of the endothelial cell layer in an engineered vessel (Scale BAR = 10 m.) (H) Immunoperoxidase staining for platelet endothelial cell adhesion molecule (PECAM) antigen shows an endothelial monolayer on the vessel lumen. (×1,000 before 33% reduction.) Reprinted with permission from Niklason LE, Gao J, Abbott WM, et al. Functional arteries grown in vitro. Science 1999;vol284: 489_¯93. © 1999 American Association for the Advancement of Science.

Ø Ongoing work in Vacanti laboratory in this area involves:

1) investigation of the ability of different polymer scaffolds to support attachment and growth of vascular endothelial and smooth muscle cells

2) comparison of different dynamic seeding methods in the construction of the vascular conduit

3) investigation of various cell-marking techniques of endothelial and smooth muscle cells

4) creation of a new pulsatile flow bioreactor capable of supporting fragile constructs during cellular maturation and architectural organization. (1)

Ø Other groups are investigating the influence of mechanical stresses on vascular cells as this relates to the tissue engineering of blood vessels (Nerem RM), the use of gene therapy in vivo to modify grafts (transfection with adenovirus expressing tissue plasminogen activator)(Kuo MD), and the addition of peptide sequences to constructs to improve endothelial cell attachment (Hubbell). (1)

Large diameter conduits:

Repair of many congenital cardiac defects requires the use of conduits to establish right ventricle to pulmonary artery contunuity. The drawback of homografts or prosthetic conduits is the lack of growth potential. Multiple operations are often necessary because of obstruction secondary to calcification or tissue ingrowth. For these reasons, researchers have pursued the creation of tissue-engineered large diameter conduits. (1)

Ø Shinóka and colleagues seeded ovine arterial and venous cells onto synthetic biodegradable tubular scaffolds of polyglactin/PGA and cultured the constructs for 7 days.

Method: Ovine artery or vein segments were harvested, separated into individual cells, expanded in tissue culture, and seeded onto synthetic biodegradable (polyglactin/polyglycolic acid) tubular scaffolds. After 7 days of in vitro culture, the autologous cell/polymer vascular constructs were used to replace a 2 cm segment of pulmonary artery in lambs. One other control animal received an acellular polymer tube sealed with fibrin glue without autologous cells. Animals were sacrificed at intervals of 11 to 24 weeks after echocardiographic and angiographic studies. Explanted tissue-engineered conduits were assayed for collagen (4-hydroxyproline) and calcium content, and a tissue deoxyribonucleic acid assay (bis-benzimide dye) was used to estimate number of cell nuclei as an index of tissue maturity. (7)

Results: The acellular control graft developed progressive obstruction and thrombosis. All seven tissue-engineered grafts were patent and demonstrated a nonaneurysmal increase in diameter Histologically, none of the biodegradable polymer scaffold remained in any tissue-engineered graft by 11 weeks. Collagen content in tissue-engineered grafts was 73.9% ± 8.0% of adjacent native pulmonary artery. Histologically, elastic fibers were present in the media layer of tissue-engineered vessel wall and endothelial specific factor VIII was identified on the luminal surface. Deoxyribonucleic acid assay, used to determine the cell density in the tissue, showed a progressive decrease in numbers of cell nuclei over 11 and 24 weeks, suggesting an ongoing tissue remodeling. Calcium content of tissue-engineered grafts was elevated, but no macroscopic calcification was found. (7)

As a result, living vascular grafts engineered from autologous cells and biodegradable polymers functioned well in the pulmonary circulation as a pulmonary artery replacement. They demonstrated an increase in diameter suggesting growth and development of endothelial lining and extracellular matrix, including collagen and elastic fibers. This tissue-engineering approach may ultimately allow the development of viable autologous vascular grafts for clinical use. However, homograft donor scarcity remains a significant problem that continues to limit its widespread clinical application. (7)

Once cells are attached to a three-dimensional biodegradable polymer, the resulting tissue construct can be implanted in vivo, where the cells continue to grow and develop a predesigned structure. While cellular structure and matrix develop, the polymer degrades ultimately leaving only the engineered tissue without foreign material. The remaining cells and matrix have the potential to remodel and organize into functional tissue that can be used for reconstructive or transplantation operation. This approach of creating tissue from autologous cells offers many potential advantages. It eliminates the problems of rejection and donor organ scarcity. This research result demonstrated evidence of functional endothelial cells and viable fibroblasts in these implanted conduits. (3)

Ø Stock and associates seeded ovine venous cells onto porous P4HB patches. After further culture of the construct, six autologously seeded patches were implanted into the proximal pulmonary artery in a patch augmentation procedure. Echocardiography showed no dilatation or stenosis, and histology showed the formation of organized tissue. Biochemical assays revealed increasing collagen, elastin, and proteoglycans content. Stock and associates also seeded ovine arterial cells onto a copolymer of PGA and PHA and used these constructs to replace 3- to 4-cm segments of the abdominal aorta in lambs. All acellular control conduits became occluded, whereas the tissue-engineered conduits remained patent. The collagen and DNA content of the constructs approached the native aorta over time, as did the mechanical strain-stress curve (1).

Ø Recently, Shin'oka and colleagues moved to the next level and implanted a tissue-engineered pulmonary artery into a human being. The patient was a 4-year-old girl who was born with a single right ventricle and pulmonary atresia. She had undergone pulmonary artery angioplasty and a Fontan procedure, but an angiogram revealed occlusion of her right intermediate pulmonary artery the following year. The group harvested a 2-cm segment of peripheral vein, isolated the cells, and expanded the cells in culture. A tubular scaffold of polycaprolactone-PLA copolymer reinforced with woven PGA was seeded with 12 million cells and transplanted after 10 days in culture. (11)

Figure 9 The Tissue-Engineering Technique. Venous-wall cells were isolated and expanded in vitro and seeded on a biodegradable polymer scaffold. The construct of cells and polymer was implanted as autologous tissue. (N Engl J Med. Vol344 No7, 532-533 www.nejm.org)

Ten days after seeding, the graft was transplanted. The occluded pulmonary artery was reconstructed with the tissue-engineered vessel graft. No postoperative complications occurred. On follow-up angiography, the transplanted vessel was noted to be completely patent (Fig. 9). Seven months after implantation, the patient was doing well, with no evidence of graft occlusion or aneurysmal changes on chest radiography. This was the first reported human implant of a tissue-engineered blood vessel constructed from cells and polymer, truly an exciting accomplishment. (11)

REFERENCES:

1. Julie R. Fuchs, MD, Boris A. Nasseri, MD, & Joseph P. Vacanti. Tissue engineering: A 21st Century Solution to Surgical Reconstruction. Ann Thorac Surg 2001 Vol 72; (577-591)

2. L’Heureux N, Paquet S, Labbe R, Germain L, Auger FA. A completely tissue engineered human blood vessel. FASEB Journal 1998; Vol 12; (47-56)

3. Principles of Tissue Engineering edited by Robert Lanza, Robert Langer and William Chick 1997; (349-359)

4. Niklason LE, Gao J, Abbott WM, et al. Functional arteries grown in vitro. Science Vol 284; (489-493)

5. Weinberg CB & Bell E. A blood vessel model constructed from collagen and cultured vascular cells. Science 1986 Vol 231; (397-400)

6. Niklason LE, Langer RS. Advances in tissue engineering of blood vessels and other tissues. Transplant Immunology 1997 Vol 5; (303-306)

7. Shinoka T, Shum-Tim D, Ma PX, et al. Creation of viable pulmonary artery autografts through tissue engineering The Journal of Thoracic and Cardiovascular Surgery Vol 115; (536-546)

8. Teebken OE, Bader A, Steinhoff G, Haverich A. Tissue engineering of vascular grafts: human cell seeding of decellularised porcine matrix. Eur J Vasc Endovasc Surg Vol 19; 2000; (381-386)

9. RM Nerem. Tissue engineering in the cardiovascular system. 12th conference of the European Society of Biomechanics,Dublin, 2000 ;(8-10)

10. Ratcliffe A. Tissue engineering of vascular grafts. Matrix Biology 2000 Vol 19 (353-357)

11. Shinoka T, Imai Y, Ikada Y. Transplantation of a tissue engineered pulmonary artery. N Engl J Med 2001 Vol 344, No 7; (532-533)