cell adhesion on - materials technology · 1990]. each leaflet roughly consists of two parts; the...

Cell Adhesion onAlginate Scaffolds forthe Tissue Engineering ofan Aortic Valve – A ReviewL. MulderReport: BMTE 02.40

Coaching: F. P. T. BaaijensC. V. C. BoutenM. I. van Lieshout

Eindhoven University of Technology,Faculty of Biomedical Engineering,October, 2002.

Cell adhesion on alginate scaffolds for tissue engineering of an aortic valve – a review

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ContentsIntroduction 3

1 The aortic heart valve 41.1 Anatomy and geometry . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4

1.1.1 Anatomy of the aortic valve . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 51.1.2 Geometry of the aortic valve . . . . . . . . . . . . . . . . . . . . . . . . . . . . 6

1.2 Mechanical characteristics of the aortic valve . . . . . . . . . . . . . . . . . . . . . 71.3 Aortic valve replacements . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8

1.3.1 Mechanical replacements, xenografts, and homografts . . . . . . . 8

2 Cell adhesion 102.1 Adhesion receptor families . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 10

2.1.1 Integrins . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 102.1.2 Cadherins, Ig-CAMs, and selectins . . . . . . . . . . . . . . . . . . . . . . . 12

2.2 Components of focal contacts . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 122.2.1 Cytoskeletal proteins . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 132.2.2 Cytoplasmic integrin-binding proteins . . . . . . . . . . . . . . . . . . . . . . 13

2.3 Signal transduction by integrins . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 142.3.1 Step 1: Signal transduction by activation of tyrosine kinases . . . . 15

2.3.1.1 The Focal Adhesion Kinase (FAK) pathway . . . . . . . . . . . 15 2.3.1.2 The Shc pathway . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 16

2.3.2 Step 2: Signal transduction by activation of the mitogen-activatedProtein kinase cascade . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 16

2.4 Other aspects of integrin signaling . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 172.4.1 Signaling by specific integrins . . . . . . . . . . . . . . . . . . . . . . . . . . . . 172.4.2 Cross-talk between integrins . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 17

2.5 Control of cell cycle, apoptosis, and gene expression . . . . . . . . . . . . . . . . 18 2.5.1 Cell cycle control . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 182.5.2 Control of programmed cell death . . . . . . . . . . . . . . . . . . . . . . . . . 192.5.3 Control of cell shape, growth, and survival . . . . . . . . . . . . . . . . . . 192.5.4 Regulation of gene expression . . . . . . . . . . . . . . . . . . . . . . . . . . . 19

2.6 RGD as cell recognition site . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 202.6.1 A tripeptide recognition sequence . . . . . . . . . . . . . . . . . . . . . . . . . 202.6.2 RGD and integrins . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 20

3 Alginate hydrogels 223.1 Alginate . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 22

3.1.1 Alginate structure and properties . . . . . . . . . . . . . . . . . . . . . . . . . 223.1.2 Alginate applications . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 25

3.2 Alginate as tissue engineering scaffold . . . . . . . . . . . . . . . . . . . . . . . . . . . 263.2.1 Mechanical properties of alginate hydrogels . . . . . . . . . . . . . . . . . 263.2.2 In vitro cell studies on unmodified alginate . . . . . . . . . . . . . . . . . . 273.2.3 In vitro cell studied on RGD-modified substrates . . . . . . . . . . . . . 283.2.4 Porous alginate scaffolds . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 31

3.2.4.1 Lyophilization . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 32 3.2.4.2 Porous alginate scaffolds . . . . . . . . . . . . . . . . . . . . . . . . . 33

3.2.5 Combining alginate with fibrin gel . . . . . . . . . . . . . . . . . . . . . . . . . 34

4 Tissue engineering of heart valves 364.1 Tissue engineering . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 364.2 Research on the tissue engineering of heart valves . . . . . . . . . . . . . . . . . 37

5 Discussion and future research 40

References 42


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Introduction

The human heart beats approximately 103,000 times every day. It is easy to imagine thatsome structures in the heart are subjected to extensive wear. One of these structures is theaortic valve, which is located between the left ventricle of the heart and the aorta. Sometimes,the aortic valve becomes unable to perform its natural function, for example when a stenosisis present. This prevents a good flow of blood through the valve. Other problems that canarise are aortic insufficiency, when the valve is unable to close properly and blood flows backinto the left ventricle, or congenital malformations of the valve. The most common surgicalprocedure to deal with these problems is valve replacement. However, there are serious limitsto the success of these replacements. These replacements are unable to grow and adapt tochanging circumstances and often patients suffer from lifelong anti-coagulation therapy. Theanatomy and geometry of the aortic valve, as well as heart valve replacements areextensively described in chapter one. Tissue engineering could be the solution to theproblems that arise with artificial heart valves; the patient’s own cells are seeded onto ascaffold material in the shape of for instance a heart valve and cultured in vitro until theconstruct is ready to be implanted into the patient’s body. Cells adhere to the scaffold materialthrough focal adhesion contacts. In these contacts, members of the adhesion receptor family,especially integrins, bind to specific sequences in the molecules of the scaffold material.Adhesion is critical for the cell to control its cell cycle and to control cell shape, growth,survival, apoptosis, and gene expression. A special recognition site for integrins is the amino-acid sequence Arginine-Glycine-Aspartic acid (RGD). This tripeptide or a larger proteincontaining the tripeptide sequence can be covalently linked to molecules of scaffold materialwhich by itself is not able to interacts with cells. One of these non-interacting scaffoldmaterials is alginate, which is derived from natural brown algae. Alginate is a copolymer andis able to form a gel by the binding of divalent cations (chapter two). Different sources ofalginate as well as the concentration of alginate and the divalent cations can greatly influencegel strength, porosity, homogeneity, and other properties. A lot of research is already done onalginate and alginate modified with RGD (containing) peptides. Chapter three describes thestudies done on alginate; from the mechanical properties to cell adhesion studies. A tissueengineering scaffold should have certain properties, one of these properties is porosity. Thereare different ways in which a material can be made porous, but this report only describes theprocess of lyophilization. Lyophilization is carried out by the freezing of the sample andsubsequently by sublimating the ice crystals which acts as porogens. Chapter four deals withthe research done on the tissue engineering of complete valve constructs, developed inpulsatile bioreactors.


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The aortic heart valve

In order to create a tissue engineered valve that resembles a natural aortic valve, it isnecessary to understand the anatomy, geometry, and function of the natural valve. Thischapter deals with the general physiology of the natural aortic valve and will give a shortsummary of diseases and congenital defects, which may lead to the replacement of the valve.In order for a tissue engineered heart valve replacement to be successful, it needs to havethe same mechanical properties as the natural valve. Therefore, the mechanicalcharacteristics of the natural aortic valve will be dealt with. Replacement of cardiac valves willbe briefly discussed; these include mechanical heart valve replacements, xenografts, andhomografts. Finally, a short review of progress in the tissue engineering of the heart valve willbe given.

1.1 Anatomy and geometryThe aortic valve is situated between the left ventricle of the heart and the aorta and preventsbackflow of blood from the aorta into the left ventricle during the diastole of the cardiac cycle.During a day, the aortic valve opens and closes 103,000 times on average, meaningapproximately 3.7 billion times in a normal life span [Thubrikar, 1990]. The position of theaortic valve is illustrated in figure 1.1 in a longitudinal cross-section of the heart and in figure1.2 the position is illustrated in a horizontal cross-section.

Figure 1.1: The position of the aortic valve. The valve is located between the left ventricle and the aorta.Adapted from http://www.cts.usc.edu.


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Figure 1.2: The position of the aortic valve in a cranial view of the heart. Adapted fromhttp://www.medfacts.com.

1.1.1 Anatomy of the aortic valve

The aortic valve consists of three leaflets and three sinus cavities. The leaflets are the mostmobile part of the entire valve and the sinuses are cavities situated behind each leaflet in thebase of the aorta. Two of the three sinuses contain apertures of the coronary arteries, whichsupply the heart muscle with blood (see figure 1.3A). Accordingly, the sinuses are namedright coronary sinus, left coronary sinus, and noncoronary (or posterior) sinus [Thubrikar,1990]. Each leaflet roughly consists of two parts; the first part of the leaflet, that coapts withanother leaflet when the valve is closed, is called the lanula. The remainder of the leafletsurface, not making contact with adjacent leaflets is referred to as the load-bearing surface ofthe leaflet or leaflet belly. The part of the lanula that bears no load is called the coaptation orredundant surface (see figure 1.3B). The only free boarder of a leaflet is called the free edgewith a thickening in the center of each leaflet, the nodulus of Arantius. The line of attachmentof the leaflets to the aorta forms a U-shaped arch and is referred to as the aortic ring. Thetops of the U-shaped arches, where the lanula of adjacent leaflets merge into the aortic ring,are called commissures. [Thubrikar, 1990; Sauren, 1981].

Figure 1.3: II: Schematic presentation of the aortic valve opened with a longitudinal incision. L, R, and Nrepresent the left, right, and noncoronary sinus, respectively. IV: Drawing of a single leaflet with the freeedge (FE), the redundant or coaptation surface (R), the line of coaptation (C), the line of attachment (A),the load-bearing surface (L). The circumferential (CD) and the radial (RD) directions are shown.Adapted from Thubrikar, 1990.

Connective tissue bundles are macroscopically visible in the leaflets of the aortic valve. Thesebundles originate at the commissures and run circumferentially like the free edge margin ofthe leaflet. The bundles form an interwoven network of fine fibers towards the center of theleaflets. Perpendicular to these bundles runs another set of fibers. These fibers are also


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macroscopically visible and anchor the middle portion of the leaflet to the aortic wall. Thefibers diverge into the aortic wall to form the U-shaped arches in each sinus as part of theaortic ring (see figure 1.4) [Sauren, 1981].

Figure 1.4: The typical collagen bundle structure in a porcine aortic valve leaflet. Adapted from Sauren,1981.

The aortic valve leaflet consists of three layers with highly specialized cells and extracellularmatrix. Two primary layers are called the fibrosa as the top layer and spongiosa as the middlelayer (see figure 1.5). The fibrosa is made up out of densely circumferentially arrangedcollagen fibers. The spongiosa is composed of loosely arranged collagen fibers and anabundant amount of glycosamino-glycans. The fibrosa and spongiosa are covered by anextension of the ventricular endocardium, called the ventricularis. This layer consists mainly ofelastin with radially aligned elastic fibers. The fibrosa and ventricularis are preloaded by theirattachment to each other, the fibrosa under compression and the ventricularis under tension[Thubrikar, 1990; Mol, 2001].

Figure 1.5: Schematic cross-sections of the aortic valve, showing the three layers of the aortic valveleaflet; the fibrosa, spongiosa, and ventricularis. Adapted from http://www.lerner.ccf.org and Mol, 2001.

1.1.2 Geometry of the aortic valve

Various dimensions of the aortic valve are shown in figure 1.6. The dimensions noted are theradius of the bases (Rb), radius of the commissures (Rc), the valve height (H), the angle ofthe free edge to the plane through the three commissures (φ), the angle of the bottom surfaceof the leaflet of the leaflet to the plane through the three commissures (α), the coaptationheight at the center (Cc), commissural height (Hs), the length of the leaflet free edge (Lf), thelength of the leaflet in the radial direction (LR), sinus depth (ds), and sinus height (hs).


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Figure 1.6: Schematic drawing of the aortic valve showing the side view of one leaflet. Radius of thebase (Rb) 11.3 to 14 mm, radius of the commissures (Rc) 8.5 to 14 mm, valve height (H) 15.7 to 19.8mm, bottom surface angle of the leaflet (α) 15 to 27 degrees, free edge angle of the leaflet (φ) 25 to 37degrees, height of the commissures (Hs) 8 to 9.9 mm, sinus height (hs) 19.8 to 28.2 mm, radius of theoutermost wall of the sinus (ds) 16.5 to 23.9 mm and coaptation height (Cc) 3.8 to 4.8 mm. Imageadapted from Thubrikar, 1990 and human aortic valve dimensions adapted from Mol, 2001.

1.2 Mechanical characteristics of the aortic valveIn order to adjust the mechanical properties of an aortic valve tissue engineering scaffold tothe physiological situation, one needs to have an understanding of the mechanical propertiesof the aortic valve leaflet. The composition of the extracellular matrix plays an important rolein the mechanical behavior of the leaflet tissue. The matrix of an aortic valve leaflet consistsof collagen, elastin, and proteoglycans. The collagen is the component that has the greatestinfluence on the mechanical strength. There are three types of collagen that can be identifiedin the leaflet; collagen type I is the most predominant (74%), followed by type III (24%) andtype V is also present in the matrix. The elastin is mainly organized as a complex network andprovides the elasticity of the valve leaflet. Proteoglycans form a gel-like substance that makesup the rest of the extracellular matrix volume. The mechanical behavior of this complex canbe described by the mechanical behavior of general soft tissues. Soft biological tissues arelike the aortic valve mainly composed of collagen and elastin fibers. The elastin fibers areable to extend up to 100% without damage to their structure. The collagen fibers however,can only extend as much as 2-4%. The general shape of a load-elongation curve obtainedfrom a uniaxial tensile test is given in figure 1.7 [Sauren, 1981].

Figure 1.7: Left: Characteristic load-elongation curve for soft biological tissues in uniaxial tensile testingat constant elongation rate. Right: the stress-strain curve derived from it. Adapted from Sauren, 1981.


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In the load-elongation curve four different phases can be identified. In the first phase, theelastin phase, the tissue offers little resistance to elongation. The only structure that transmitsforce is the elastin network. In the second phase, the transition phase, more and morecollagen fibers become aligned and uncoiled and take up more and more of the forcetransmission. In the third phase, all of the collagen fibers are recruited and transmit all theforce. In the final phase, the collagen fibers start to get damaged and eventually this will leadto the complete rupture of the tissue.This force (load)-elongation curve can be converged into a stress-strain curve (see figure1.7). The stress is calculated from the force by dividing it by the cross-sectional area of thesample. The strain is obtained by dividing the elongation by the original length of the sample.From this stress-strain curve a high and low modulus of elasticity can be calculated by takingthe tangent of the curve at a certain point. The low modulus of elasticity refers to the elastinphase and the high modulus of elasticity to the collagen phase. Sauren (1981) tabulated thetensile properties of a human aortic valve leaflet and found that the leaflet exhibits ananisotropic behavior. Table 1.1 below shows the values found by Sauren et al. (1981).

Radial CircumferentialEl (MPa) 2.76x10-3 – 1.12x10-2 1.99x10-2

Eh (MPa) 0.17 – 2.27 3.53 – 5.98Table 1.1: The high and low modulus of elasticity of aortic valve tissue in the radial and circumferentialdirection. Adapted from Sauren, 1981.

1.3 Aortic valve replacementsWhen an aortic valve is unable to perform its natural function due to disease ormalformations, it may be necessary to replace the natural valve with a prosthesis. When theaortic valve is stenotic it offers resistance to the blood flow out of the heart. Aortic stenosis ismost commonly caused by the calcification of the leaflets of the aortic valve. A congenitallyunicuspid, bicuspid, or qaudricuspid aortic valve is more prone to such calcifications. Variationin the leaflet size of one aortic valve also increases the chance of calcification. Other originsof stenosis are postinflammatory calcification of rheumatic origin and cuspid fibrosis. Anotherdisease of the aortic valve is aortic insufficiency. In this case the aortic valve allows blood toflow back into the left chamber of the heart. Aortic insufficiency can occur either by itself or inassociation with aortic stenosis. Insufficiency is generally caused by postinflammatorydisease of rheumatic origin, aortic root dilatation, congenital defects, or infective endocarditisand leads to regurgitation of blood into the left chamber of the heart [Thubrikar, 1990]. The most common treatment for the diseases described above is the replacement of thenatural valve with an artificial valve. These artificial valves include mechanical replacements,xenografts, and homografts (see next section). These options all have specific advantagesand disadvantages as will be described in the following section. To avoid the disadvantagesof the different artificial replacements, a new method has been developed; tissue engineering,which is outlined in the final section of this chapter.

1.3.1 Mechanical replacements, xenografts, and homografts

Nowadays, mechanical replacements of the aortic valve are the mostly used valvereplacement. Generally speaking there are three different classes of mechanical valves(figure 1.8). The ball and cage mechanical heart valve: the blood pressure in the left chamberof the heart pushes the ball upwards so the blood can pass through the valve. Second, thetilting disc mechanical valves and third, the bileaflet mechanical valves. The advantage ofmechanical heart valves is their good durability, although damage and wear resulting in wearmarks and leaflet escape are well known. Mechanical heart valve prostheses patients sufferfrom lifelong anticoagulation therapy combined with a high risk of bleeding and blood damage(e.g. hemolysis, thromboembolic complications) [http://www.hia.rwth-aachen.de].


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Figure 1.8: The three types of mechanical heart valves; ball and cage, tilting disc, and bileafletmechanical valve. Adapted from http://www.leva.leeds.ac.uk, http://www.me.utexas.edu, andhttp://www.onevalveforlive.com respectively.

The most commonly used xenografts are bovine or porcine pericardial valves. The pericardialtissue is treated with gluteraldehyde for fixation and attached to a frame. These valves havethe advantage that they are abundantly available, have low risk for thrombosis and thereforeanticoagulation treatment is not necessary. However, these xenografts have low durabilityand thus show high degradation and rate of failure [Mol, 2001].

Homografts, also known as human aortic allografts, are the heart valve replacements thatresemble the natural aortic heart valve the most. They are obtained from human cadaversand can be preserved at -196°C. The quality of the valves depends on the details of thefreezing and thawing protocols, the time between death of the donor and harvest andadditional warm and cold ischemic intervals [Schoen, 1999]. These valves are not chemicallycrosslinked with gluteraldehyde and therefore have better mechanical properties, which givesthem a longer lasting time. They are not thrombogenic and have low risk of infection. Theirdisadvantage lies in the fact that they are only limited available and can cause an immuneresponse.

Nowadays, the most common procedure for end-stage valvular heart disease is valvereplacement. The problems that occur with the mechanical, xenograft, and homograftreplacements have led to a complete new and innovative solution: tissue engineering. Intissue engineering the patient’s own tissue is used, so no problems with rejection will occurand the tissue will be able to grow and adapt to changing circumstances. The tissueengineering process and the current research done on heart valves will be elaborated inchapter 4.


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Chapter 2

Cells adhere to their surrounding extracellular matrix (ECM) through cell-surface receptors.Cell adhesion is critical for the genesis and maintenance of both three-dimensional structureand normal function in tissues and plays essential roles in cell motility, cell proliferation, celldifferentiation, regulation of gene expression, cell survival, and embryonic morphogenesis[Retta, 1999]. Cell adhesion comprises a cascade of four steps: initial cell attachment, cellspreading, organization of an actin cytoskeleton, and formation of focal adhesions [LeBaron2000]. In these focal adhesions, the cell-surface receptors are transmembrane glycoproteinsand connect cytoskeleton components like actin filaments, microtubules, and intermediatefilaments to the extracellular matrix molecules, which are usually fibrillar in nature [Petit, 2000;Aplin, 1998]. The major transmembrane components present in focal adhesions are part of alarge family called integrins. Cell adhesion complexes are dynamic structures, which are ableto process mechanical and chemical signals from the extracellular environment, which willlead to an intracellular reaction. The opposite is also true: the functions of cell adhesioncomplexes are regulated by biochemical events within the cells [Giancotti, 1999]. This chapterwill review the different components of adhesive junctions and the way in which their signaltransduction pathways work. Also, the role of the integrins in cell cycle control, programmedcell death, and gene expression will be elaborated upon. Finally, the most commonrecognition sequence in extracellular matrix proteins and its role in tissue engineering will bediscussed.

2.1 Adhesion receptor families When looking at cell to cell adhesion and cell to matrix adhesion, there are four families of celladhesion receptors: integrins, cadherins, immunoglobulin-cell adhesion molecules (Ig-CAMs),and selectins. In this section, the structure and role of Ig-CAMs, cadherins, and selectins willbe discussed briefly. The integrins will be elaborated upon in more detail, because they arethe most common and well-understood component of cell adhesion structures.

2.1.1 Integrins

The integrins are a family of cell-surface glycoproteins that act as receptors for ECM proteins,or for membrane-bound counter receptors on other cells. Most integrins are expressed on awide variety of cells, and most cells express several integrins. All integrins are heterodimerscomposed of two subunits, α and β, and each αβ combination has its own binding specificityand signaling properties [Giancotti, 1999; Hynes, 1992]. The α-subunits vary in size between120 and 180 kDa and are each noncovalently associated with a β-subunit (90-110 kDa). Eachsubunit consists of a large extracellular domain, a single membrane spanning region, and ashort cytoplasmic domain (figure 2.1). To this day the integrin family includes at least sixteendistinct α subunits and eight β subunits which can associate to form more than twenty distinctintegrins [Aplin, 1998]. Each integrin recognizes several ECM matrix proteins and a singleECM protein can be recognized by different integrins. Table 2.1 is a summary of most of theknown integrins and the ligands or counter receptors that they bind to. The binding ofintegrins to their ligands depends on the presence of extracellular divalent cations (Ca2+ orMg2+, depending on the integrin), reflecting the three or four divalent-cation-binding domainsin the large extracellular part of the α chain [Alberts, 1994]. Integrins differ from other cell-surface receptors in that they bind their ligands with a lowaffinity and that they are usually present in 10-100 fold higher concentrations on the cellsurface.


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Figure 2.1: Schematic representation of the two subunits of an integrin receptor. On the α chain(left), four divalent cation-binding sites are represented. Also notice the long extracellular part,the transmembrane section, and the short cytoplasmic tail. (Adapted from Alberts (1994)).

Subunits Ligands and counterreceptors Binding siteβ1 α1 Collagens, lamininβ1 α2 Collagens, laminin DGEAβ1 α3 Fibronectin, laminin, collagens RGDβ1 α4 Fibronectin, VCAM-1 EILDVβ1 α5 Fibronectin (RGD) RGDβ1 α6 Lamininβ1 α7 Lamininβ1 α8 ?β1 αv Vitronectin, fibronectin RGDβ2 αL ICAM-1, ICAM-2β2 αM Fibrinogen, factor X, ICAM-1β2 αX Fibrinogen GPRPβ3 αIIb Fibrinogen, fibronectin, von Willebrand factor, vitronectin,

thrombospondin RGD,KQAGDV

β3 αv Vitronectin, fibrinogen, von Willebrand factor, thrombospondin,fibronectin, osteospondin, collagen

RGD

β4 α6 Laminin ?β5 αv Vitronectin RGDβ6 αv Fibronectin RGDβ7 α4 Fibronectin, VCAM-1 EILDVβ8 αv ?

Table 2.1: An overview of the different integrins and their ligands and the bindingsite on the ligand.(Adapted from Hynes (1992)).

The integrins however can only bind their ligands when they exceed a certain minimal numberof integrins at one spot, called a focal contact and hemidesmosomes. So when the integrinsare diffusely distributed over the cell surface, no adhesion will occur, but when after a certainstimulus these integrins cluster in focal contacts their combined weak affinities give rise to aspot on the cell surface which has enough adhesive (sticking) capacity to adhere to theextracellular matrix. This is a very useful mechanism, because in this way cells can bindsimultaneously but weakly to large numbers of matrix molecules and still have the opportunityto explore their environment without losing all attachment to it by building or breaking downfocal contacts. If the receptors were to bind strongly to their ligands, cells would probably beirreversibly bound to the matrix, depriving it from motility. This problem does not arise whenattachment depends on multiple weak adhesions.


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2.1.2 Cadherins, Ig-CAMs, and selectins

The cadherins form a transmembrane protein family that shares an extracellular domain,which is made up out of multiple repeats of a cadherin-specific motif. Cadherins can be foundin specialized sites of cell-to-cell adhesion named adherence junctions. In these junctions,cadherins can establish links with actin filaments of the cytoskeleton. The structure of acadherin consists of an external domain with an amino terminal, which repeats five times.Further, there is a single transmembrane section, and a cytoplasmic domain with a carboxyterminal. The binding site of the cadherins is localized in the last repeating unit of theextracellular domain, which is the only repeat that is not bridged by calcium binding sites. Thecalcium bridges give rigidity to the receptor molecule [Aplin, 1998]. Cadherins play animportant role in tissue organization and maintenance. They directly interact with actinfilaments through α-actinin, an actin-bundling protein, and specialized intracellular proteins,called catenins.

Proteins of the Immunoglobulin-Cell Adhesion Molecule (Ig-CAM) family consist of one ormore copies of the Ig fold, a compact structure with two cysteine residues arranged as twoantiparallel β sheets. Ig family adhesion receptors have a large amino-terminal extracellulardomain, a single transmembrane helical segment, and a cytoplasmic tail.One of the most important functions of the Ig-CAMs lies in the guidance of axons and in theestablishment and maintenance of neural connections in the development of the nervoussystem. Another function of the Ig-CAMs lies in the immune system, where they areexpressed by T-lymphocytes. These receptors play important roles in antigen recognition,cytotoxic T-cell functions, and lymphocyte recirculation. Other Ig-CAM family adhesionreceptors are located on vascular endothelial cells and play an important role in leukocytetrafficking towards inflamed tissues. There is little known about the interaction of Ig-CAMswith protein of the cytoskeleton [Aplin, 1998].

The selectins are a small family of lectin-like adhesion receptors composed of threemembers, L-, E-, and P-selectin. The structure of selectin on the extracellular side consists ofan amino-terminal domain, followed by an epidermal growth factor (EGF)-type domain, andtwo to nine complement regulatory protein repeats. Furthermore, a transmembrane helicalsegment, and a short cytoplasmic tail make up the rest of the receptor. The role of selectinslies in the leukocyte adherence to endothelial cells and to platelets during inflammation. P-selectin is present in latent form in endothelial cells and platelets and is rapidly expressed ontheir surfaces through translocation from secretory granules, when they are activated bythrombin or other agonists. There is much less known about signal transduction by selectinsthan is the case for integrins, cadherins, and Ig-CAMs [Aplin, 1998].

2.2 Components of focal contacts

The term cell adhesion describes all forms of cellular communication due to direct contactbetween cells or between cells and the extracellular matrix [Löster, 2000]. This cellularcommunication takes place in adhesive junctions called focal contacts. At focal contacts,clusters of integrins bind externally to ECM proteins and internally to cytoplasmic proteins thatin turn bind to the cytoskeleton (see figure 2.4). It is believed that focal contacts are dynamicstructures that change in size as the cell adhesion process progresses. As the focal contactmatures, the adhered actin filaments extend and bundle to form structures called stressfibers. [Aplin, 1998]. This paragraph describes the different cytoskeletal proteins that areinvolved in the connection between integrins and the cytoskeleton. After that, the cytoplasmicproteins that interact with integrins in order to establish inside-out signaling are discussed.

2.2.1 Cytoskeletal proteins

The cytoplasmic structural proteins of the focal contact that directly bind to integrins includetalin and α-actinin, which in turn bind to other cytoplasmic proteins. These proteins aresummarized in table 2.2, and elaborated a little further here. Figure 2.2 is a schematic


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representation of the cytoplasmic structural proteins and their interactions with integrins andvarious other proteins.Talin: Talin is a homodimer of two 270 kDa polypeptides arranged in an antiparallelorientation. The N-terminal region interacts with phospholipids in membranes, whereas the C-terminal portion interacts with sites on the β1 or β3 integrin cytoplasmic domain. Talin has arole in connecting the actin proteins to focal adhesions and it is crucial in the initial formationof new focal adhesions.α-Actinin: α-Actinin is an actin-bundling protein that comprises two identical polypeptides,each of which is approximately 104 kDa in size. The protein can be divided into three regions:the amino-terminal actin binding site, a central region with four α-helical motifs, and a carboxyterminal region. α-Actinin has been shown to interact with the cytoplasmic domains of theintegrin β1, β2, and β3 subunits. α-Actinin is necessary for the association of actin moleculesto focal adhesions.Tensin: This protein was so named because of its proposed role in connecting actin to focaladhesion and maintaining mechanical tension. Tensin consists of two 200 kDa polypeptidechains and has three actin binding sites per chain. In addition, tensin has been reported tobind to vinculin, paxillin, Src, FAK, PI-3K, and p130CAS (see below). Vinculin: Vinculin is a 116 kDa polypeptide. Vinculin binds to actin, α-actinin, talin, andpaxillin. It has globular head and a rod-like tail region. The first 120 amino acids in the headregion are essential for binding talin and α-actinin, whereas paxillin an actin bind to the rod-like tail region.Paxillin: Paxillin is a 68 kDa protein and seems tethered to the membrane at focal adhesions.Paxillin binds to a region of 140 amino acid residues in the carboxy terminus of FAK, and ithas been hypothesized that this interaction might coordinate the signals from focal adhesionto the cytoplasm or cytoskeleton.Filamin: Filamin is an actin-binding protein that is involved in the formation of actin networks.Filamin exists as a homodimer with the two polypeptides associating only at the carboxyterminal region, whereas the amino terminal ends link with actin. The association of filaminwith integrins might be important in spreading and extension of lamellopodia during cellmovement and in phagocytosis [Aplin, 1998].

Figure 2.2: This diagram illustrates the patterns of binding interactions of focal contact structuralproteins including talin (Tal), vinculin (Vinc), α-Actinin (α-Act), tensin (Tens), and paxillin (Pax). Adaptedfrom Aplin, 1998.

2.2.2 Cytoplasmic integrin-binding proteins

The cytoplasmic domains of integrins are involved in bi-directional transfer of informationacross the membrane. Integrin cytoplasmic tails interact with several proteins that mayparticipate in integrin-mediated functions including signal transduction. Figure 2.3 sheds somelight on the many protein interactions involved in integrin signaling. Table 2.2 summarizesthese proteins in a fashionable order and shows the associated integrin subunits.


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Intracellular proteins suggested to interact directly with integrinsProtein Associated integrin

Cytoskeletal proteins Talin β1; β1A, β1D; αIIbβ3α-Actinin β1; β2; β3Filamin β2; β1A, β1D, β7Tensin -Paxillin β1Vinculin -

Intracellular signaling proteins Cytohesin-1 β2

FAK β1ILK β1β3-endonexin β3ICAP-1 β1ARack1 β1, β2, β5CIB αIIbCalreticulin α tails; α6β1Calnexin β1, α6

Table 2.2: Intracellular proteins suggested to interact directly with integrins. FAK = Focal AdhesionKinase, ILK = Integrin Linked Kinase, ICAP-1 = Integrin Cytoplasmic-domain-Associated Protein-1,Rack1 = Receptor for activated protein kinase C, CIB = Calcium- and Integrin-Binding protein [Adaptedfrom Hemler, 1998; supplemented from Aplin, 1998].

Figure 2.3: This diagram recapitulates information concerning proteins that directly bind to integrins.Adapted from Aplin, 1998.

2.3 Signal transduction by integrinsIntegrins can signal through the cell membrane in either direction: the extracellular bindingactivity of integrins is regulated from the inside of the cell (inside-out signaling), while thebinding to the ECM elicits signals that are transmitted into the cell (outside-in signaling).Integrin signaling and assembly of the cytoskeleton are closely linked. As the integrins bind tothe ECM, they become clustered and associate with a cytoskeletal and signaling complex hatpromotes the assembly of actin filaments. The reorganization of actin filaments into largerstress fibers, in turn, causes more integrin clustering, thus enhancing the matrix binding andthe organization by integrins in a positive feedback system. As a result, ECM proteins,integrins, and cytoskeletal proteins assemble into aggregates on each side of the membrane,called focal contacts. A schematic representation of focal contact assembly can be seen infigure 2.4 [Giancotti, 1999; Bray, 1998].


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Figure 2.4: Matrix binding promotes integrin clustering and association with the cytoskeleton. This inturn promotes further integrin clustering and matrix organization in a positive feedback system. RGD =Arg-Gly-Asp integrin-binding motif; Tal = talin; Pax = paxillin; Vin = vinculin; CAS = p130CAS. Adaptedfrom Giancotti, 1999.

The development of a mature focal contact assembly and its associated signaling complexesrequires integrin aggregation, integrin occupancy, tyrosine kinase activity, and actin integrityin a coordinated and hierarchical manner [Aplin, 1998].

2.3.1 Step 1: Signal transduction by activation of tyrosine kinases

The overall goal of the signaling pathways of integrins is initiation of the MAP-kinase cascade.The MAPK pathway will be explained in the next section. First, there are two ways that lead tothe MAPK cascade and they both involve tyrosine activation. Activation of tyrosine kinases isan important event for integrin-mediated signal transduction. In many cells FAK and Src arethe kinases most directly responsive for integrin-dependent cell adhesion. However, some β1and αv integrins also activate the tyrosine kinase Fyn and, through it, the adapter protein Shc.Both pathways will be discussed here:

2.3.1.1 The Focal Adhesion Kinase (FAK) pathway

The FAK pathway is activated by most integrins. FAK interacts directly or through thecytoskeletal proteins talin, tensin, and paxillin with the cytoplasmic tail of integrin β subunits.Upon integrin-mediated cell adhesion FAK is rapidly phophorylated, which forms a bindingsite for the Src homology 2 (SH2) domain of Src or Fyn. The Src kinase then phosphorylatesa number of focal adhesion components. The major targets include paxillin, tensin, andp130CAS (CAS), a protein that recruits the adapter proteins Crk and Nck. After this Srcphosphorylates FAK at Tyr925, creating a binding site for the complex of the adapter Grb2 andRAS guanosine 5’-triphosphate exchange factor mSOS. These interactions link FAK to thesignaling pathways that modify the cytoskeleton and activate mitogen-activated protein kinase(MAPK) cascades, also known as extracellular-signal-regulated kinases (ERK) cascades[Giancotti, 1999; Aplin, 1998]. The JNK cascade is also activated by adhesion, which is acascade that is very similar to the MAPK cascades. Figure 2.5A groups all the players of theFAK pathway and shows their interactions.


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2.3.1.2 The Shc pathway

In this pathway, caveolin-1 interacts with an integrin, which creates a binding site for Src-family kinases, such as Fyn, Yes, and Lck. Upon integrin binding to ECM, Fyn becomesactivated, and its SH3 domain interact with a proline-rich site in Shc. Shc is thenphosphorylated by Fyn at Tyr317 and combines with the Grb2-mSOS complex. Theseinteractions lead to the activation of MAP (ERK) kinase cascades. Although most integrinsinteract with caveolin-1 and Fyn, only a few of the integrins can activate Fyn and therebyrecruit Shc. In figure 2.5B this pathway is schematically shown.

Figure 2.5: Model of the (A) FAK and (B) Shc pathways. N = NH2-terminal, C = COOH-terminal, Y397 =Tyr397, P = phosphotyrosine, Y = tyrosine, Cav = caveolin-1. Adapted from Giancotti, 1999.

2.3.2 Step 2: Signal transduction by activation of the mitogen-activatedprotein kinase cascade

Between the step from RAS to MAPK in the previous section, some intermediate steps arerequired. RAS activates Raf-1, which in turn activates MEK. MEK phosphorylates MAPKleading to direct activation of the MAPK cascades. The functions of the MAPK cascades arestill very obscure, but a few possible functions will be highlighted. Once a MAP-kinase is activated, it relays signals downstream by phosphorylating variousproteins in the cell, including other protein kinases and gene regulatory proteins. For instance,MAP kinases may phosphorylate the Jun protein, which combines with the newly made Fosprotein to form an active gene regulatory protein called AP-1. The AP-1 protein then turns onadditional genes [Alberts, 1994]. Another example is the phosphorylation and activation ofpp90RSK (a.k.a. MAPKAPK-1). Once activated pp90RSK can move into the nucleus andcontribute to the regulation of transcription by AP-1 [Aplin, 1998]. Another candidate for integrin mediated MAPK activation is the microtubule network. MAPK isassociated physically with and has enzymatic activity toward microtubule components,implicating that microtubules and microtubule-associated MAPK activity plays a role in theregulation of integrin mediated signaling events and the formation of focal adhesions andstress fibers.Inhibition of MAPK activity suppresses cell migration, but not cell adhesion or in situ cellspreading, suggesting that integrin mediated MAPK activation is not required for these initialevents in this system [Aplin, 1998].Integrin mediated MAPK activity may persist at 60%-100% of maximum levels for periodsranging from 15 minutes to two hours or more. Thus, the time course of integrin mediatedMAPK activation is likely to vary, based on the type of cell, integrin, and matrix involved[Aplin, 1998; Aplin, 1999].


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2.4 Other aspects of integrin signaling

2.4.1 Signaling by specific integrins

The FAK activation and activation of the MAPK cascade, can be initiated by several differentintegrins. It seems likely that individual integrins participate in specific signaling events. Table2.3 tabulates several examples of specific biological actions of individual members of theintegrin family that cannot be substituted by other members of the family.

Specific integrin Biological effectβ1c Growth inhibitionβ6 Growth promotionαvβ3 Control of apoptosis in endothelial cellsα6β4 Growth control in epithelial cellsα5β1 Control of apoptosisα2β1 Growth control in smooth muscle cellsα6β1 Differentiation of skeletal muscle cellsαv integrins Cellular calcium increase

Table 2.3: Specific actions of individual integrins. Adapted from Aplin, 1998.

2.4.2 Cross-talk between integrins

Signaling of integrins to each other usually takes the form of one integrin regulating the abilityof another integrin family member to support cell adhesion, motility, or phagocytosis. Forinstance, the MAPK cascade triggered by one integrin can modulate the ligand binding affinityof another. Also, several integrins can share interactions with transmembrane or cytoplasmicproteins, which could be another mechanism for cross talk. Table 2.4 gives a few examples ofcross talk between different integrins.

Adhesion receptors involved Biological effectα5β1/αvβ3 α5β1 affects motility mediated by αvβ3αvβ3/α5β1 αvβ3 affects phagocytosis mediated by α5β1αMβ2/FcIII αMβ2 and the FcIII receptor cooperate in the

respiratory system αvβ5/uPAR A protease receptor modulates αvβ5-mediated

motilityL-selectin/β2 L-selectin activates β2 integrinsPECAM-1/β2 and β1 The Ig-family receptor PECAM-1 activates β2

and β1 integrinsN-cadherin/β1 and β3 Integrins activate N-cadherinαIIb/β3/other β1 integrins Trans-inhibition of one integrin by another

Table 2.4: Cross talk between integrins or with other adhesion receptors. Adapted from Aplin, 1998.


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2.5 Control of cell cycle, apoptosis, and gene expressionAdhesion is important in the control of cell growth and division. Integrins activate growth-promoting signaling pathways that are responsible for the anchorage requirement. So cellsneed adhesion through integrins to complete a cell cycle. Loss of attachment to the matrixcauses apoptosis in many cell types. Cell adhesion also has important consequences interms of the regulation of gene expression during development of tissues. The followingparagraphs will deal with several important properties of cells and the influence of adhesion inthe control of cell cycle, apoptosis (or anoikis), and gene expression.

2.5.1 Cell cycle control

The molecular machinery that dictates movement of a cell through the phases of the cell cyclecomprises a series of cyclins and CDKs. Levels of expression of various cyclin proteinschange during different phases of the cycle. Cyclin A/B-CDC2 complexes are required forentry into the M phase, cyclin A-CDK2 complexes for transit of the S phase, whereas cyclinD-CDK4/CDK6 complexes are associated with G1, and cyclin E-CDK2 with the G1/Stransition. A point late in the G1 phase is called the restriction point and once cells have pastthis point they no longer require adhesion to a substratum to complete the cell cycle.Movement through the restriction point is associated with activation of cyclin D-CDK4,6complexes and of cyclin E-CDK2 complexes. These two complexes phosphorylate twomembers of one family, namely Rb and p107. Upon phosphorylation of Rb by activatedcyclin-CDK complexes, E2F is released, which can transcriptionally activate genes that arerequired for further progress through the cell cycle (see figure 2.6). There is evidence that MAPKs can increase expression of cyclin D1. Cyclin D1 expression ispositively regulated by p42/p44 MAPK and negatively regulated by the related p38 kinasepathway. By this means early signaling events in the MAPK pathway and related pathwayscan have an impact on the G1 cell cycle transition. The biochemical basis for cell cycle arrestseems to be an inhibition of cyclin E-CDK2 kinase activity because of increased levels of thep21 and p27 CDIs and is thus similar to the effect of loss of anchorage. This suggests thatintegrin mediated focal contact formation supports cell cycle traverse and that disruption orreduction of focal contacts promotes arrest. [Aplin, 1998]. Cell attachment through integrins may also facilitate exit from the cell cycle and providesignals for differentiation. For example, the α6β1 integrin may in part promote exit from thecell cycle, because the cytoplasmic domain of α6 inhibits paxillin signaling [Giancotti, 1999].Other integrins are unable to efficiently activate Shc or FAK or lack the ability to cooperatewith growth factor receptors, which may facilitate differentiation and induce exit from the cellcycle.

Figure 2.6: Anchorage regulation of the cell cycle. Key regulators of the cell cycle including cyclins,cyclin-dependent kinases (CDKs), the p21 and p27 inhibitor proteins, and the Rb protein are shown atthe approximate times of their major periods of activity. Positive (+) or negative (-) influences of integrinmediated cell anchorage on these cell cycle regulators are shown. Adapted from Aplin, 1998.


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2.5.2 Control of programmed cell death

Loss of attachment to the matrix causes apoptosis in many cell types. This phenomenon iscalled anoikis and may help to maintain the integrity of tissues; it would prevent cells thathave lost contact with their surroundings from establishing themselves at inappropriatelocations. FAK appears to play a major role in conveying signals from the ECM. Because FAKbinds PI 3-kinase, the protective effect against anoikis may be the result of PI 3-kinase-mediated activation of protein kinase B/Akt. Akt promotes survival by phosphorylating andthereby inactivating two pro-apoptotic proteins, Bad and caspase-9. FAK-expressing cellsthat are growth arrested because of a loss of matrix adhesion accumulate phosphorylated Rb,apparently because of the inactivity of cyclin D-CDK4,6 and cyclin E-CDK2. Inhibiting theactivity of the Rb target E2-F transcription factor protects cells against apoptosis, suggestingthat phosphorylated Rb can provide an apoptotic signal. Like cell growth, the ECM can controlanoikis in an integrin-specific manner. Certain integrins induce expression of anti-apoptoticproteins and therefore integrin mediated attachment to ECM is a general requirement for cellsurvival, but survival under special circumstances may require a particular integrin.Because most cells in adult organisms are not actively dividing, it is likely that other cellsurface proteins, such as cadherins, override the growth-promoting, but not the survival-promoting, effects of integrin and growth factor receptors.Anoikis is likely to be important in the maintenance of tissue architecture, as it would ensurethe demise of cells that detach from their original site in tissue [Giancotti, 1999].

2.5.3 Control of cell shape, growth, and survival

When cells come in contact with the ECM, their usual response is to extend filopodia,apparently to sample the terrain. Integrins at the tip of the filopodia bind to the ECM andinitiate the formation of focal adhesions. Actin-rich lamellipodia are then generated, oftenbetween filopodia, as the cell spreads on the ECM. Fully developed focal adhesions andassociated actin stress fibers originate. Integrins regulate cell spreading and migrationthrough activation of the Rho-family of small guanine nucleotide-binding proteins. Each ofthese Rho-related proteins controls the actin cytoskeleton by interacting with multipledownstream effectors. The FAK-Src complex also plays a role in cell migration, perhaps bypromoting the disassembly of focal adhesions at the trailing edge of the cell. Thus, the ECMand the cytoskeleton are interdependent, and the geometry and physical properties of theECM available to the cell are important for cell spreading and motility. The coordinated controlof cell shape, survival, and growth by integrins is likely to be important in the establishmentand maintenance of tissue architecture [Giancotti, 1999; Howe, 1998].

2.5.4 Regulation of gene expression

Regulation of genes by integrins is dependent on the integrin species involved in adhesionand therefore on the proteins involved in the MAPK cascade. Table 2.5 shows integrins andthe gene they regulate.

Receptor involved Gene regulated Cell typeβ1 integrins IE genes Human monocytesβ1 integrins Tissue factor Human monocytesβ1 integrins Milk genes Mouse breast cellsβ1 integrins Muscle-specific genes Quail myoblastsβ1 integrins Metalloproteases Synovial fibroblasts β1 integrins Cytokines Chondrocytesβ2 integrins E-selectin Endothelial cellsαvβ3 Metalloproteases Melanoma cellsα5β1 IE genes Colon carcinoma cellsα5β1 Bcl-2 CHO cellsVarious integrins IL-2 T cells

Table 2.5: Regulation of gene expression by cell adhesion receptors. Adapted from Aplin, 1998.


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2.6 RGD as cell recognition siteExtracellular matrices are made up out of an insoluble network of protein and carbohydratethat is laid down by cells and that fills most of the intercellular spaces. Matrices in differentlocations in the body consist of different combinations of collagens, proteoglycans, elastin,hyaluronic acid, and various glycoproteins such as fibronectin and laminin. Virtually all of theextracellular matrix glycoproteins and collagens that have been identified interact with cells,and much of the control of cellular behavior appears to originate in response to theseinteractions. The proteins that contain the Arginine-glycine-aspartic acid (Arg-Gly-Asp orRGD) sequence among their amino acid array, together with the integrins that serve asreceptors for them, constitute a major recognition system for cell adhesion. The RGDsequence is critical for the interaction of the ECM proteins with the adhesion receptors[Ruoslahti, 1987; Ruoslahti, 1996].

2.6.1 A tripeptide recognition sequence

The RGD cell adhesion sequence was first discovered in fibronectin in 1984. An incompletelist of adhesion proteins with RGD sites includes fibronectin, vitronectin, fibrinogen, vonWillebrand factor, thrombospondin, laminin, entactin, tenascin, osteopontin, bone sialoprotein,and, under some conditions collagen [Ruoslahti, 1996]. Other sequences can also promoteadhesion to adhesion receptors. For instance, the KGD sequence can be a highly potent andspecific binder of αIIbβ3. The glycine position can be occupied by a number of differentindividual amino acid residues, or even by two residues, with retention of integrin bindingactivity. The NGR sequence is another variation of RGD capable of binding to various RGD-directed integrins, albeit with a low affinity. In addition, the peptide SDGR was reported toinhibit cell attachment in a manner similar to RGD peptides when presented to cells insolution. Another sequence, KQAGDV, appears to be a mimic of the RGD sequence,because its binding to the αIIbβ3 integrin, for which it is essentially specific, is inhibited byRGD peptides [Ruoslahti, 1996]. Changes in the peptides as small as the exchange of alanine for the glycine or glutamic acidfor the aspartic acid, which constitute the addition of a single methyl or methylene group tothe RGD tripeptide, eliminate adhesion activities [Ruoslahti, 1987].

2.6.2 RGD and integrins

It seems likely that the sites recognized by the various integrins evolved from one primordialrecognition site and that this is why the specificities of the present-day integrins are stillclosely related. In each integrin case, the ligand-binding site is near or at the binding site fordivalent cations. The α subunit also contains one or more ligand-binding sites, just as the βsubunits do. These sites are localized in the divalent cation binding sequences. The RGDbinding is independent of divalent cations and causes the extrusion of the divalent cationsfrom the integrin. This suggests that the ligand binds to a cation binding site at the integrinand that the primary role of the adjacent binding site for divalent cations is to keep the site ina conformation receptive for the ligand. This one hypothesis on the role of the divalent cationbinding sites.

The aspartic acid may be important because of its potential to contribute to divalent cationbinding; another hypothesis regarding the binding of ligands to integrins postulates that theligand provides a coordination site for divalent cation binding. The integrin α subunits containfour sequences that resemble the EF-hand divalent cation-binding motif of other proteins. Theintegrin sequences, however, lack one of the conserved acidic amino acids that serves as acoordination site in EF-hands. The suggestion is that the aspartic acid of RGD and the otheradhesion motifs that contain an aspartic acid residue may provide the missing coordinationsite when an integrin binds to its ligand. The amino acid residue on the carboxy-terminal siteof the RGD is more significant in integrin binding than the amino-terminal one [Ruoslahti,1996].


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At least eight, and possibly as many as twelve, of the currently known twenty or so integrinsrecognize the RGD sequence in their ligands. Table 2.6 states some integrin-bindingsequence motifs and the integrins recognizing them.

KQAGDV LDV/IDS RLD/KRLDGS

RGD L/IET R…D YYGDLR/FYFDLR

(α5β1) (α5β1) α5β1α8β1αvβ1

αvβ3 αvβ3αvβ5αvβ6αvβ8

αIIbβ3 αIIbβ3(α2β1) α2β1(α3β1)(α4β1)(α7β1)

α4β1αMβ2

αLβ2α1β1

Table 2.6: Various integrin target sequences and the integrins that bind to them. Binding that is weak orthat is only seen under special conditions is indicated by parenthesis. Adapted from Ruoslahti, 1996.

The situation with the RGD sequence is complex, because the same tripeptide is recognizedby several receptors and because these receptors can recognize the RGD sequence in oneprotein but not in another. One explanation for this could be that the RGD sequence servesas a shared binding site, whereas the specificity is generated by a second binding site uniqueto each protein ligand. Alternatively, the specificity could reside in the conformation of theRGD tripeptide, and the role of surrounding sequences would be to force the RGDdeterminant into an appropriate conformation. It may be that the RGD sequences of proteinswith incidental cell attachment activity happen to be in the conformation of one of theadhesion protein sequences. An inactive RGD sequence, on the other hand, may either notbe available at the surface of the molecule containing it, or, if available, its conformation maynot fit any of the receptors [Ruoslahti, 1987; Ruoslahti, 1986].

Adhesive peptides can be used in two different ways: When attached to a surface, theypromote cell attachment, whereas when presented in solution, they prevent attachment thatwould otherwise occur. Both modes of using the peptide have found applications: Surfacescoated with RGD peptides are being investigated for improvement of tissue compatibility ofvarious implantable devices, and soluble peptides targeted for individual integrins are apromise as potential drugs for the treatment of a number of diseases, like diabetes.Farthest along among the RGD applications is the use of RGD peptides, or compoundsdesigned to mimic RGD, as anti-thrombotics that act by inhibiting the function of the αIIbβ3integrin. Other applications that are being explored include the targeting of the αvβ3 integrinin osteoporosis. Osteoclasts attach to bone through this integrin, and inhibiting theirattachment with peptides prevents bone degradation. A new development is the finding thatαvβ3 is selectively expressed on endothelial cells that are engaged in angiogenesis. RGDpeptides and anti-integrin antibodies can prevent tumor growth by interfering with theangiogenesis that the growing tumor needs to maintain its blood supply. RGD peptides canalso affect tumor cells in a more direct fashion; peptides coinjected with tumor cells into theblood stream can prevent the subsequent growth of tumor nodules in tissues [Ruoslahti,1996].


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Chapter 3

As mentioned in chapter one, tissue engineering of tissues requires a three-dimensionalscaffold. These scaffolds are designed to deliver the cells to the desired site in the body,provide a space for new tissue formation, and potentially control the structure and function ofthe engineered tissue. Hydrogels and other scaffold materials for tissue engineering mustmeet a number of design criteria to function appropriately and promote new tissue formation.These criteria include classical parameters like structure and mechanics, as well as biologicalperformance parameters like suitability for cell adhesion [Lee, 2001]. In this chapter the basicstructure and the applications of the hydrogel alginate will first be discussed. After that thecriteria which alginate should satisfy as a tissue engineering scaffold will be discussed. Theseinclude the mechanical properties, porosity, and homogeneity. Finally, a review of in vitro andin vivo studies that were performed using alginate will be given.

3.1 AlginateAlginate is a collective term for a family of polymers with a wide range in chemicalcomposition, sequential structure, molecular size and hence in functional properties. Thepresent paragraph will deal with how features such as porosity, swelling behavior, long timestability, and gel strength depend on the chemical structure and the molecular size of thealginate molecule.

3.1.1 Alginate structure and properties

Alginate’s occurrence in nature is mainly limited to the marine brown algae (Phaeophyta).Alginate exists in the brown algae, as the most abundant polysaccharide comprising up to40% of its dry matter. It is located in the intercellular matrix as a gel containing sodium,calcium, magnesium, strontium, and barium ions. Its main function is believed to be skeletal,giving both strength and flexibility to the algal tissue. In molecular terms alginate is a binary copolymer of 1-4 linked β-D-mannuronic acid (M) andα-L-guluronic acid (G), illustrated in figures 3.1 and 3.2. The monomers are arranged in ablock-wise pattern along the chain with homopolymeric regions of M and G termed M- and G-blocks respectively. These M- and G-blocks are interspaced with regions of alternatingstructure (MG-blocks). Alginates are able to form gels by the binding of divalent cations to theG-units. The size of a cooperative unit can be as large as twenty monomers. The so-called“egg-box” model explains this characteristic property (figure 3.3). The affinity for variousdivalent cations is Pb2+ > Cu2+ > Cd2+ > Ba2+ > Sr2+ > Ca2+ > Co2+, Ni2+, Zn2+ > Mn2+. Table 3.1summarizes the origin and composition of the most common types of alginates [Thu, 1996].

Figure 3.1: The monomer composition in alginate. G: α-L-guluronate, M: β-D-mannuronate.


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Figure 3.2: the alginate chain.

Source FG FM FGG FMM FMG FGM

Laminaria japonica 0.35 0.65 0.18 0.48 0.17 0.17L. digitata 0.41 0.59 0.25 0.43 0.16 0.16L. hyperborea Blade 0.55 0.45 0.38 0.28 0.17 0.17 Stipe 0.68 0.32 0.56 0.20 0.12 0.12 Outer cortex 0.75 0.25 0.66 0.16 0.09 0.09Lessonia nigrescens 0.38 0.62 0.19 0.43 0.19 0.19Ecklonia maxima 0.45 0.55 0.22 0.32 0.32 0.32Macrocystis peryfera 0.39 0.61 0.16 0.38 0.23 0.23Durvillea antarctica 0.29 0.71 0.15 0.57 0.14 0.14Ascophyllum nodosum Fruiting body 0.10 0.90 0.04 0.84 0.06 0.06 Old tissue 0.36 0.64 0.16 0.44 0.20 0.20

Table 3.1: Origin and composition (sequence frequency (F)) of alginates. Adapted from Thu, 1996.

Gel strength: In general, alginates rich in guluronate residues form strong, brittle gels, whileM-rich alginates form softer, more elastic gels. The elastic modulus of an alginate geldepends on the number and the strength of the crosslinks and the length and stiffness of thechains between crosslinks [Thu, 1996; Draget, 1997]. The increased gel strength withincreasing G content (and increasing length of G-blocks) stems from association of long G-blocks and shortening of the elastic segments [Draget, 1997]. The modulus also depends onthe crosslinking ions; the gel strength increases with the affinity between polymer andcrosslinking ions. The compressive modulus increases with increased amount of gelling ions,presumably due to increased crosslinking density, but at a certain concentration ofcrosslinking ions the moduli of the gels begin to decrease. This is known as the inflectionpoint [Lee, 2000]. Alginate hydrogels lose more than 60% of their initial mechanical strength within 15 hours ofexposure to physiological buffers. This can be attributed to the loss of divalent ions from thehydrogels by ion exchange with monovalent ions in the surrounding fluid [Lee, 2000]. Thisproblem can be avoided by adding calcium ions to the surrounding fluid.

Stability: For many uses, it is important to be aware of the factors that determine and limit thestability of alginate solutions and the chemical reactions responsible for the degradation. Themolecular weight of alginate in solution may be severely reduced over a short period of timeunder conditions favoring degradation. The saccharide linkages are susceptible to both acidicand alkaline degradation and oxidation by free radicals. The degradation is at its minimumaround neutrality and increases in either direction. The increased instability at pH values less


24

than 5 can be attributed to proton catalyzed hydrolysis, whereas the reaction responsible forthe degradation at pH 10 and above is β-alkoxy-elimination [Thu, 1996; Draget, 1997].

Swelling and shrinkage: Viscosity is mainly affected by the molecular weight/size. Below acritical value for molecular weight, depending on the concentration, the gel-forming abilitiesare reduced. Above a critical value the mechanical properties and swelling are governed bythe monomeric composition, and the block-structure. By increasing the content of G, the gelsbecome mechanically stronger, and the stability in the presence of anti-gelling ions (Na+,Mg2+) is enhanced. High-G gels exhibit high porosity, low shrinkage during gel-formation, anddo not swell after drying. By increasing the M content the gels become softer and moreelastic; they shrink more and their porosity is reduced. They swell after drying and are easierto break down in the presence of EDTA or phosphate ions than high-G material [Thu, 1996].

Porosity: Diffusion characteristics are essential for the use of alginate gels as tissueengineering scaffolds. It is therefore important to know the pore sizes and pore sizedistribution. The highest diffusion rates of proteins, indicating the most open pore structure,are found in gels made from high-G alginates. An explanation for this behavior is that high-Ggels, with their long G-blocks and their short elastic segments become more of a stiff openand static network compared to the more dynamic and entangled network structure of thelow-G gels with their relative long elastic segments [Draget, 1997]. A model is given in figure3.3 [Thu, 1996].

Figure 3.3: A model of alginate gel network in gels made from alginates with differentcomposition. Notice the “egg-box” shape of the G-blocks due to binding with divalent cations.Adapted from Thu, 1996.

Gelling kinetics: Alginate gels are often regarded as non-equilibrium gels. After a cation hasinduced random binding of G-blocks, the resulting “egg-box” structure does not necessarilycorrespond to the most stable conformation [Thu, 1996]. Gelation rate increases withtemperature. GDL (see gel homogeneity) probably hydrolyzes more quickly at highertemperatures, releasing calcium ions from CaCO3 more rapidly [Kuo, 2001].

Gel homogeneity: Alginate gel beads prepared by the dialysis method (see next section)often exhibit a concentration inhomogeneity in that the polymer concentration is considerablylower in the center of the bead than at the edges. Alginate diffuses from the center of the geltowards the gelling zone, leading to a depletion of alginate in the center [Thu, 1996]. Thisgelling zone thus moves from the surface towards the center of the gel. This is explained bythe activity of alginate, which will equal zero in this zone, and alginate molecules will diffusefrom the internal, non-gelled part of the gelling body towards the zero activity region resultingin a depletion of alginate in the center part of the gel [Draget, 1997]. Due to the very rapid andirreversible binding of Ca2+ to the G-blocks, it is impossible to obtain a homogeneous gel bysimply adding Ca2+ to an alginate solution. The most common way to overcome this problemis to introduce Ca2+ into the alginate solution in an inactive form with a subsequent controlledrelease and gel formation; the so-called internal gelation. This can be achieved by addingcalcium-carbonate and D-glucono-δ-lactone (GDL). With this system, the time-delayedrelease of crosslinking calcium ions allows the calcium-alginate suspension, with or withoutcells, to be molded into a complex geometry before gelation occurs [Kuo, 2001]. Maximum


25

inhomogeneity is obtained with a low molecular weight alginate in a solution containing a lowconcentration of the gelling ion and in the absence of non-gelling ions. Maximumhomogeneity is reached by a high molecular weight alginate gelled with high concentrationsof both gelling and non-gelling ions. [Thu, 1996; Draget, 1997]. Structural uniformity in tissueengineering scaffolds is necessary not only for uniform cell distribution, but also for well-controlled material properties. Uniform pore size and distribution ensure diffusion of nutrientsinto all areas of the gel and the removal of metabolic wastes from the system [Kuo, 2001].Homogeneity of alginate gel samples can be examined by measuring the difference in wet todry weight ratios of subsequent slices of that sample.

Sterilization: Because of their natural origin, alginates have to be cleared of impurities andmicroorganisms before their use as a tissue engineering scaffold. Autoclaving an alginatedispersion causes a decrease in viscosity of about 64% [Vandenbossche, 1993]. The degreeof breakdown was related to the temperature and to the time of exposure. Three heatingstages at lower temperatures than autoclaving (120-135°C) may also cause a breakdown ofthe alginate chains and may not deliver a sterile product.The exposure time and temperature during ethylene oxide sterilization causes a decrease inviscosity. A possible solution can be found in ethylene oxide sterilization when using lowtemperatures [Vandenbossche, 1993].γ-Irradiation is often disastrous. It is generally believed that under these conditions, oxygenmolecules are rapidly depleted with formation of the very reactive hydroxy free radicals[Draget, 1997].After membrane filtration, no significant changes in viscosity and molecular weight areobserved. For applications in which the molecular weight is critical, membrane filtration is themethod of choice [Vandenbossche, 1993].

3.1.2 Alginate applications

Because of their biocompatibility, abundance in source, and low prices, alginates have beenwidely used in the food industry [Kuo, 2001] as thickeners and emulsifying agents, and in thecosmetics industry.They have also been processed into gel beads, encapsulating living cells as a means ofimmunoprotection. These gel beads are also used in drug delivery applications, such asmicroencapsulation of pancreatic islets and alginate-based bioreactors for large-scalemanufacture of biological products, for example, monoclonal antibodies. Alginate has alsobeen used as a bio-artificial matrix for cartilage generation and fundamental studies onentrapped chondrocytes [LeRoux, 1999; Draget, 1997]. Furthermore, alginate is widely usedfor wound dressings [Lee, 2001; Suzuki, 1998].

Immobilization of cells by entrapping them in a hydrogel is generally carried out by mixing thecells with a water-soluble polymer, and subsequent gelling of the polymer by addingcrosslinking agents.

Figure 3.4: The microencapsulation method: A solution of alginate with cells is slowly drippedinto another solution, containing gelling ions. Alginate gel beads are formed containing thecells within.


26

For alginate, adding cations such as calcium or strontium induces gelling. By dripping thealginate-cell mixture into a solution containing divalent cations, the droplets willinstantaneously form gel-spheres by gelation, entrapping the cells within a three-dimensionallattice of ionically crosslinked polymer (figure 3.4) [Thu, 1996].

Gel entrapment in alginate, as the technique is known today, has some limitations due to boththe inherent nature of the alginate molecule itself, as a biodegradable material, and to thenature of the gel as a reversible ionic network. As a consequence of the latter, substanceswith high affinity for calcium ions such as phosphate or citrate will sequester the crosslinkingcalcium ions and consequently destabilize the gel [Thu, 1996]. This fact will cause anincreased swelling potential due to an increased osmotic pressure and less elastic strengthinside the liquid capsule, making it difficult to obtain a prolonged mechanical stability [Draget,1997]. Since the calcium ions can be exchanged with other cations, the gel will also bedestabilized by high concentrations of non-gelling ions, such as sodium and magnesium [Thu,1996].

3.2 Alginate as tissue engineering scaffoldAlginate lacks cellular interaction. That is, cells are unable to adhere to the alginate network.As was mentioned in chapter 2, this has large consequences for the cell growth,differentiation, gene expression, and proliferation of most cell types. Alginate however still is apromising material for a scaffold in tissue engineering of an aortic heart valve. Rowley et al.(1999) showed that cellular interaction could be established by modifying alginate with apeptide containing the RGD recognition sequence. As told in chapter 2, this is critical for theinteraction of the ECM molecules (alginate in this case) with the adhesion receptors on thecell surface. The disadvantage of alginate is its degradation characteristics. Degradationthrough the loss of divalent cations into the surrounding medium is an uncontrollable andunpredictable process, even with dissolved calcium ions in the surrounding medium. Studiesare being performed to covalently crosslink alginate with various types of molecules anddifferent crosslinking densities to precisely control the mechanical, swelling, and degradationproperties of alginate [Lee, 2001]. In this section, literature will be reviewed on the studiesperformed to determine the mechanical properties of alginate. The cell studies that wereperformed on plain alginate and RGD-modified alginate will be reviewed. After that, themethods of creating porous scaffolds are discussed and special attention will be given to thelyophilization process. Finally, the combination of alginate with fibrin gel is discussed.

3.2.1 Mechanical properties of alginate hydrogels

The mechanical properties of alginate gels are important, because as a tissue engineeringscaffold it must be resistant to the same mechanical loads as the original tissue, in this casethe aortic valve leaflets. Different testing methods can be applied to measure certainmechanical properties of alginate gels. Stokke et al. (2000) studied calcium alginate gelsthrough rheologic measurements to determine the relations between chemical compositionand concentration of the alginate. They used alginates with low (39%), intermediate (50%),and high (68%) fractions of G residues originating from different brown algae. Dynamicviscoelastic characterization of the alginate samples was carried out by determining thestorage and loss moduli, G’ (ω) and G” (ω) at T = 20°C and ω = 6.28 rad s-1. Thus, the storageand loss modulus were determined while using three different G residue contents, twodifferent calcium sources were examined and different concentrations of calcium were used. Itwas shown that there was a significant increase in both moduli when high-G alginates werecompared with intermediate-G and low-G alginates. Increasing calcium concentration resultedin an increase in both moduli. There was no significant difference in moduli between the twocalcium sources that were used (CaCO3 and CaEGTA). Lee and Mooney (2000) investigatedthe possibility that the mechanical properties of alginate hydrogels could be better controlledby covalently linking the alginate molecules with various molecules with different size andstructure, including adipic dihydrazide, lysine, and poly- (ethylene glycol)- diamines. Shearmoduli (G) of the alginate hydrogels were obtained. The reaction with each of the bifunctional


27

crosslinking molecules led to the rapid formation of hydrogels, and hydrogels with widelyvarying mechanical properties were formed with this method. As the amount of crosslinkingmolecules increased, the shear modulus of the hydrogels increased likewise likely due to theincreasing number of interchain crosslinks. It is generally considered that interchain crosslinksmainly contribute to the mechanical strength of the network. However, above a certainconcentration of crosslinking molecules the moduli of the gels began to decrease. Theconcentration at which this occurs is known as the inflection point. At this point the nature ofthe crosslinking molecules affects the properties of the hydrogels. It was shown that themechanical properties and swelling of alginate hydrogels can be tightly regulated by usingdifferent kinds of crosslinking densities. The mechanical properties of the hydrogels aremainly controlled by the crosslinking density but are also dependent on the type ofcrosslinking molecules: alginate gels created with adipic dihydrazide showed the highestshear modulus and when the gels were crosslinked with calcium, the shear modulus waslowest. Kuo et al (2001) performed an investigation in which the compressive modulus andstrength of three-dimensional alginate samples were measured. They compared two differentgelling formulas; CaCO3-GDL and CaSO4-CaCO3-GDL. It was shown that the compressivemodulus and strength of gels made with the CaCO3-GDL system were superior to those ofgels made with the CaSO4 system. The slower gelation rate of the CaCO3-GDL systemprovided time for the CaCO3 microparticles to evenly disperse throughout the suspensionduring mixing before complete gelation occurred. It was also shown that the compressivemodulus and strength of the gels increased with increasing calcium content and increasingalginate concentration. Improvement in mechanical properties with increasing alginateconcentration was attributed to the increase in polymer chain density and entanglement.LeRoux et al (1999) determined the equilibrium and viscoelastic properties of alginate gelscrosslinked with calcium ions as a function of alginate concentration and duration of exposureto physiological saline solution. Compressive and shear stress relaxation tests and oscillatoryshear tests were performed to measure the stress and strain at two time periods after storagein saline compared to no saline exposure. The effect of concentration was determined bytesting 1-3 % alginate gels in a bath of physiological saline and calcium chloride. After 15hours of exposure to saline, the compressive, equilibrium shear, and dynamic shear modulidecreased by 63, 84, and 90% of control values, respectively. The material propertiesexhibited no further changes after 7 days of exposure to NaCl. All moduli significantlyincreased with increasing alginate concentration. The observed decrease in compressive andshear stiffness for alginate gel after exposure to Na+ was significant and indicated thatphysiological conditions will soften the gel over a time period up to 7 days after gelation.

3.2.2 In vitro cell studies on unmodified alginate

A lot of studies have been performed regarding the cellular behavior in or on alginatehydrogels. Some studies concerned unmodified alginate gels and others studied alginatemodified with the RGD sequence. The following section deals with the cell studies performedon these two cases in literature. An overview of the studies done on plain, unmodified alginateis given in table 3.2.

Scaffold Cells Seeding Method Results Reference3.18%alginate gels

MC3T3-E1osteoblasts

6.7 x 106 cells/ml seeded bymixing the cells with thealginate solution prior togelation.

Successful incorporation and uniformdistribution. Cells were sphericallyshaped and did not adhere to thescaffold.

Kuo et al. (2001)

1 and 2% al-ginatesolutions forgelling

C2C12mouseskeletalmyoblasts

25000 cells/cm2 onto alginatedisks

No attachment of cells to the surface.Cells maintained spherical shape.

Rowley et al.(1999)

Alginatehydrogel

MC3T3-E1, RCOosteoblasts

10000 cells/cm2 onto alginatesamples

Showed little adherence tounmodified alginate

Alsberg et al.(2001)

Alginatehydrogel

Calf Chon-drocytes

50000 cells per well on 24-wells plate onto alginatediscs

5% of the chondrocytes attached tothe surface of the gels

Genes et al.(2001)

2% alginatehydrogelculture discs

Calf Chon-drocytes

150 x 106 cells/ml on alginatediscs

No apparent attachment to thepolymer

Genes et al.(2000)


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3-D Alginatesponges

hepatocytes 1 x 106 cells/sponge, cellsolution was absorbed by thehydrophilic alginate sponges

Cells were efficiently entrappedwithin sponges. No proliferation, butmaintained metabolic activity for 1week. Toxicity tests indicated thatcells maintained viability. Cells werespherical in shape and formedaggregates.

Glicklis et al.(2000)

Alginate discs hepatocytes 30000 cells/cm2 onto alginatediscs

Minimal adhesion efficiency andreduced viability due to the lack ofadherence.


Alginatebeads

Perichon-drium cells

4 x 106 cells/ml of gel by theencapsulation method.

Little proliferation of cells Van Osch et al.(2000)

2% ultrapurelow viscositymannuronicacid alginate

Schwanncells

4 x 106 cells/ml seeded bymixing the cells with thealginate solution prior togelation.

Cells remained viable and no sign ofinfection. Cells showed neuritesprouting.

Mosahebi et al.(2001)

0.14%alginate beads

Ovine in-tervertebraldiscnucleuspulposescells

1.0 x 104 cells/10µl bead,using the cell encapsulationmethod.

Majority of the cells retained theirviability and maintained a sphericalmorphology.

Pattison et al.(2001)

Table 3.2: Overview of studies done on unmodified alginate.

All of the studies performed in the table above were done using plain, unmodified alginatesamples. As can be seen, a lot of different cell types were tested, but all showed more or lessthe same results. When cultured in or on unmodified alginate, the cells retained their sphericalmorphology and showed little proliferation. In all cases, the cells remained viable. Kuo et al.(2001) incorporated osteoblasts in alginate samples and kept the samples in completemedium supplemented with 0.0030 M CaCl2, to maintain the size and shape of the cell-hydrogel construct. After 3 weeks a hematoxylin and eosin staining revealed that the cellswere distributed evenly over the 3-D structure and assumed a spherical shape. Thetransparent nature of alginate gels makes it possible to use light microscopy to evaluate thecells inside an alginate construct. The studies of Rowley et al. (1999) were mostly performedon the surface of alginate gels. Alsberg et al. (2001) also performed an in vivo study toobserve bone growth in alginate samples. The cells that were seeded were rat-derivedcalvarial osteoblasts (RCO) and they were incorporated in the gels. The gels were implantedsubcutaneously by injection prior to gelation in mice. After 4, 16, and 24 weeks the mice weresacrificed and the bone formation was studied. Implants with no cells were used as a negativecontrol and no bone formation was observed. However, a dramatic increase in bone formationwas found over time in all samples containing transplanted cells. Disperse, small nodules ofbony tissue could be observed surrounded by the alginate matrix at 4 weeks. Histologicalexamination confirmed the beginning of bony tissue development that resembled immaturewoven bone formation. At 16 and 24 weeks, continued bone apposition was observed.Osteoblasts were observed to line newly formed bone surfaces, and osteocytes occupiedlacunae. This study shows the biocompatibility of alginate and its possible application astissue engineering scaffold for bone cell transplantation. The two studies mentioned in table3.2 concerning alginate beads (van Osch et al. (2000) and Pattison et al. (2001)) mainlyfocussed on culturing cells in alginate in order to do measurements on production ofextracellular matrix molecules. Van Osch et al. (2000) measured the production of smallamounts of glycosaminoglycan, but no collagen. When the serum was replaced by growthfactors, an increased glycosaminoglycan production and induction of collagen was found.This indicates that alginate by itself poorly stimulates the production of extracellular matrixmolecules by cells. This is in line with the theory from chapter 2; cell adhesion to thesurrounding matrix has influence on the production of ECM molecules. Mosahebi et al. (2001)demonstrated that neurite growth in alginate hydrogel was reduced to 43% when compared toscaffolds that didn’t contain alginate.

3.2.3 In vitro cell studies on RGD-modified substrates

In order to stimulate cell adhesion to alginate gels, they can be modified by covalently linkingthe RGD peptide sequence to the alginate molecules. This can be done by utilizing aqueous


29

carbodiimide chemistry. 1-ethyl-(dimethylaminopropyl) carbodiimide (EDC), water-solublecarbodiimide, was used to form amide linkages between amine containing molecules (likeRGD) and the carboxylate moieties on the alginate polymer backbone. Figure 3.5 gives thereaction in which the GRGDY pentapeptide is conjugated to the alginate polymer backbonethrough the terminal amine of the peptide [Rowley, 1999].

Figure 3.5: Reaction scheme of peptide coupling to alginate. Amide bond formation is mediated by thecarbodiimide through the carboxyl group of the alginate and the N-terminal amine of the GRGDYpentapeptide. Adapted form Rowley, 1999.

As can be seen in the reaction scheme above, the RGD sequence is often attached to thealginate molecules in the form of a larger peptide in which the RGD sequence is present. Areview of studies performed with cells on substrates modified with adhesion peptides iscaptured in table 3.3.

Scaffold cells Adhesionpeptide

Results Reference

1% alginate C2C12skeletalmyoblasts25000 cells/cm2

GRGDYthroughout 3-D structure.34 nmol/l

Myoblasts attached and began spreading by 4h, andcell number increased by day 3. Number of cellsper gel reached a maximum at 72 hrs. After thatcell number decreased, due to fusion into multi-nucleated myofibrils.

Rowley et al.(1999)

2% alginate C2C12skeletalmyoblasts25000 cells/cm2

GRGDY onsurface of gel.1 nmol/cm2

Myoblasts attached and began spreading by 4h, andcell number increased by day 3. Number of cellsper gel reached a maximum at 72 hrs. After thatcell number decreased, due to fusion into multi-nucleated myofibrils.

Rowley et al.(1999)

2% high-Galginate

C2C12skeletalmyoblasts2400cells/cm2

GGGGRGDY10 fmol/cm2

The rate of proliferation significantly increased asthe G-content of the alginate increased. The highestG-content substrates promoted extensive fusion ofthe myoblasts, with subsequently increasing levelsof muscle creatine kinase activity.

Rowley et al.(2002)

2% low-Galginate


GGGGRGDY10 fmol/cm2

The lower G-content adhesion substrates supportedlittle to no fusion. Many of the myoblasts in thenonfused cultures tended to lose adhesivity to thesubstrate, and the cultures were a mixture of spreadcells, rounded cells, and aggregated cell clumps.

Rowley et al.(2002)

2% high-Galginate


GGGGRGDYvaryingdensities

As peptide density increased, more promotion ofextensive fusion and increasing creatine kinaseactivity. Between 10 and 100 fmol/cm2 nodifferences observable.

Rowley et al.(2002)

Aminophaseglass slides

Fibroblasts,smoothmuscle cells,endothelialcells, 10000to 40000cells/ml

RGDS,YIGSR,VAPG,VGVAPG,KQAGDV,RGES asnegativecontrol. 0.5nmol/cm2

Significant differences in adhesion andproliferation rates were observed for each of thecell types evaluated, but similar trends for matrixproduction were observed. The most adhesivesurface, KQAGDV for smooth muscle cells,displayed the highest adhesion and the least matrixproduction, while the least adhesive surface RGES,had the greatest matrix production. This was alsotrue for fibroblasts and endothelial cells.

Mann et al. (1999)

Aminophaseglass slides

Fibroblasts,smoothmuscle cells,endothelialcells, 10000to 40000cells/ml

RGDS,YIGSR,VAPG,VGVAPG,KQAGDV,RGES asnegativecontrol. 2.0nmol/cm2

The higher peptide density led to a greater amountof adhesion to all of the peptides than the controlsurfaces. Less matrix production was observed thanfor the lower peptide density.

Mann et al. (1999)


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Glasssurfaces andacryloyl-PEGhydrogels

Vascularsmoothmuscle cells,40000cells/ml forglass and 1 x106 cells/mlfor gels

RGDS,VAPG,KQAGDVand RGES asnegative, 2nmol/cm2 forglass and 1.4µmol/ml

It was shown that the matrix production on a percell basis for cells grown on glass slides with 0.04pmol/ml soluble TGF-β1 in the medium, isdecreased when no TGF-β1 is present in themedium. More matrix production was observed ingels with tethered TGF-β1 than in gels withoutgrowth factor.

Mann et al. (2001)

Aminophaseglass slidesand acryloyl-PEGhydrogels

Smoothmuscle cells,10000,12000, 40000and 100000cells/ml

RGDS,VAPG,KQAGDVand RGES asnegative, 0.2nmol/cm2

and 2.0nmol/cm2

Cell migration was higher on surfaces with 0.2nmol/cm2 of adhesive ligand than on controlsurfaces, but it was lower on surfaces with 2.0nmol/cm2 of adhesive ligand than it was on controlsurfaces. Further, cell proliferation was lower onadhesive surfaces than it was on control surfaces,and it decreased as the ligand density increased.Similarly, in the peptide-grafted hydrogel scaffolds,cell proliferation was lower in scaffolds containingthe adhesive peptides than it was in controlscaffolds. After 7 days of culture, more collagenper cell was produced in control scaffolds than inscaffolds containing adhesive peptides. In addition,collagen production decreased in scaffolds as theligand concentration increased.

Mann et al. (2002)

Microtiterplates andtitanium alloysamples

Osteo-progenitorcells; humanbone marrowstromal cells(HBMSC),10000cells/cm2

RGD, RGDS,GRGD,GRGDS,GRGDSPC,GRGFSPC,YIGSR,cyclo-RGDFV,cyclo-DFVRA,cyclo-DFKRG 0-5mM

Concerning linear peptides, the results demonstratethat the short linear peptides consisting of eitherRGD or GRGD does not affect the humanosteoprogenitor cell adhesion. RGDS induced aninhibition of cell adhesion onto vitronectin. About100% of cell adhesion inhibition was observed onvitronectin using GRGDS, GRGDSPC, orGRGDSPK, while GRGFSPC did not affect cell-attachment. The cyclic peptides induced asignificant increase of cell adhesion.

Verrier et al.(2002)

Alginatehydrogel

MC3T3-E1cells, andRCO cells,10000cells/cm2

G4RGDY, 0-150 fmol/cm2

Both cell types demonstrated little adhesion tounmodified hydrogels. In contrast, both adhered toand spread on the G4RGDY-modified hydrogelsurfaces. Soluble G4RGDY peptide decreased thenumber of cells attached to the modified hydrogelsafter 4 hours in a dose-dependent manner.

Alsberg et al.(2001)

High-M(MVM) andhigh-G(MVG)alginate

C2C12skeletalmyoblasts,80000cells/cm2

G4RGDY,56-58%peptideincorpora-tion

Cells adhered to the biomaterial substrate, adproliferated and fused to form muscle fibers overtime. Confocal microscopy examination of thesecultures indicated that cells directly adherent to thehydrogel were not fusing to form myofibers, butcells at increasing distances from the gel werefusing. The calcium concentration can regulatemyoblast differentiation.

Rowley et al.(2002)

Alginatehydrogel Chondrocytes

, 50000 cellsper well in24-well plate

G4RGDY, nodensityspecified

Substrate crosslink density significantly alterschondrocyte interaction with modified alginatesurfaces. Increasing crosslinking density leads togreater chondrocyte adhesion. Chondrocyteattachment to alginate was greatly enhanced bybonding RGD ligand as evidenced by an increasednumber of cells spread onto the surfaces. Thepresence of antibodies to specific integrins reducedthe amount of attached cells.

Genes et al.(2001)

2% Alginatehydrogel Chondrocytes

, 150 x 106

cells/ml

G4RGDY, nodensityspecified

Chondrocytes cultured in alginate remainedsuspended in the material with no apparentattachment to the polymer. The presence of RGDpeptide enables cells to adhere to alginate.

Genes et al.(2000)

Table 3.3: An overview of studies with RGD modified surfaces.

The work of Rowley and Mooney focuses on the surface modification of alginate hydrogelswith RGD containing peptides. Their studies never included adherence in a three-dimensionalscaffold, but always on the surface. They found that myoblasts seeded onto GRGDY-coupledalginate hydrogels attached and began spreading within 4 hours and that the spreading wasgreatly enhanced at 24 hours. By day 3, the cell number had increased significantly whencompared to the previous days. They also studied the effects of soluble GRGDY as an


31

inhibitor of adhesion and found a strong decrease in cell adhesion to the modified alginatesurface. This showed that an RGD containing peptide has the ability to mimic extracellularmatrix molecule binding sites and stimulate adhesion to materials that are otherwise unable tointeract with cells. Rowley and Mooney also established that the alginate type and the densityof the RGD peptide can control the cell function. The rate of proliferation significantlyincreased as the G-content of the alginate increased, even though the type and density of theadhesion ligand remained constant. By varying the peptide density from 1-100 fmol/cm2, theyfound that the myoblasts spread to an increasing extent, and the rate of myoblast proliferationincreased with increasing ligand density up to 30 fmol/cm2. The rate of proliferation reached amaximum at a ligand density of 30 fmol/cm2. Recent studies by Rowley et al. (2002) showedthat the crosslinking ion density, in this case calcium cations, controls myoblast differentiation.Cells were cultured on gels formed from high mannuronic (M) acid alginates, which bind lesscalcium than the high guluronic (G) acid alginates. Quantification of calcium release fromdifferent alginate substrates confirmed an enhanced calcium release from the high M gels, ascompared to the high G gels. Myoblasts did maintain the ability to adhere and multiply on bothhigh M and high G substrates. However, myoblast fusion was greatly inhibited on the high Malginates when compared to high G alginates, presumably due to increased calciumconcentrations surrounding the high M gels. The strong adhesion of cells to modifiedsubstrates has a negative effect on the production of extracellular matrix proteins. Mann et al.(1999) demonstrated that the matrix production by smooth muscle cells decreased withincreasing surface peptide density. Later studies by Mann et al. (2002) with adhesion peptidemodified surfaces showed that not only the matrix production was inhibited by strongeradhesion, but it also leads to a decrease in proliferation and migration of smooth muscle cellsgrown on these modified surfaces. Mann et al. (2001) demonstrated that the decrease inmatrix production of cells grown on peptide modified surfaces can be compensated withgrowth factors. When transforming growth factor β1 (TGF-β1) is tethered to the scaffold, or insoluble form in the medium, a significant increase in matrix production was observed whencompared with modified surfaces without TGF-β1.Verrier et al. (2002) investigated several linear and cyclic adhesion peptides to study their rolein cell adhesion. The linear peptides that were tested were RGD, RGDS, GRGD, GRGDS,GRGDSPC, GRGFSPC, and YIGSR. The cyclic peptides were cyclo-RGDFV, cyclo-DFVRA,and cyclo-DFKRG. The effect of these adhesion peptides was studied through inhibition ofcell adhesion, by using the peptides as soluble inhibitors. They found that the short linearadhesion peptides RGD and GRGD inhibited cell adhesion up to about 40%. However, theaddition of a serine amino-acid residue to RGD or GRGD peptides (RGDS or GRGDS)strongly inhibited cell adhesion (∼90%). The other RGD containing peptide (GRGDSPC)showed the same inhibition of adhesion, whereas the non-RGD containing peptides showedno inhibition of adhesion. To study the differences between the linear and cyclic peptides,different concentrations of both types of peptides were coated onto cell culture plates. Celladhesion assays showed a dose-dependent effect of these peptide coatings. Cells boundstrongly to the cyclic peptide (∼100%) up to the coating concentration of 1 µM while a similarcell adhesion to the linear RGD peptides was obtained at a concentration of 1 mM. Thesedata suggest that cyclic adhesion peptides in particular constitute a good candidate topromote cell adhesion.

3.2.4 Porous alginate gels

In order for a tissue engineering scaffold to successfully function, it should be very porous andthe pores should have the right dimensions. The architecture of a porous scaffold can controlthe extent of scaffold vascularization and tissue ingrowth. For the ingrowth of capillaries,however not needed in the tissue engineering of heart valve leaflets, the pores of an implantshould be around 200-300 µm and highly interconnected [Zmora, 2002]. Pore morphologyand orientation can significantly affect cell seeding and nutrient and waste transport within thethree-dimensional matrix. Control over these properties of a tissue engineering scaffold canprove to be very useful. There are several ways in which a material can be made porous,these include solvent casting and particulate leaching, gas saturation, gas foaming, andlyophilization. In this section a review of studies concerning the development of porousalginate scaffolds will be given.


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3.2.4.1 Lyophilization In the lyophilization (or freeze drying) process, materials are frozen prior to the sublimation ofthe ice in a freeze drying apparatus. The growing ice crystals during freezing create the poresinside a three-dimensional sample and the subsequent sublimation of the crystals leaves athree-dimensional porous scaffold (see figure 3.6).

Figure 3.6: Schematic representation of the freeze drying process. Adapted from Kang, 1999.

Kang et al. (1999) were able to influence the pore size of three-dimensional freeze driedgelatin samples by adjusting the temperature of freezing prior to the lyophilization process.They found that as the freezing temperature was lowered, the compactness of the freezedried samples increased. At -20°C the pore size of the samples was around 250-300 µm andthe wall thickness was thin, which resulted in a weak and brittle scaffold. As the freezingtemperature was lowered to -80°C and -196°C a scaffold was created with smaller pore size(45-50 µm) and thicker walls. This resulted in a more elastic material. The difference in poresize and inner structure of the hydrogels reflect the differences in heat transfer rates duringthe freezing process of the hydrogel [Kang, 1999]. It is possible that at a higher freezingtemperature, the number of nuclei of ice crystallization initially formed is smaller than that at alower freezing temperature, leading to an increased final size of ice crystals. Since the largerice crystals push to expand the hydrogel chains to a greater extent, the pore size of hydrogelswill be increased, while the structure undergoes destruction. Rapid cooling causes formationof many nuclei of ice crystals, resulting in the formation of smaller-sized pores. In the finalstage, frozen materials are dried by sublimation of ice crystals under vacuum at atemperature below freezing.It is possible to influence the orientation of the pores by subjecting the samples to atemperature gradient. According to Kang et al. (1999) this gradient can only be created duringfast cooling methods, for example liquid nitrogen and -80°C. The lower the freezingtemperature, the more oriented the pore structure will be [Kang, 1999] (see figure 3.7).

Figure 3.7: Different degrees of orientation during freeze drying. From left to right: -196°C, -80°C, and -20°C. Adapted from Kang et al. 1999.


33

3.2.4.2 Porous alginate scaffolds

Scaffold Method In vitro /in vivo

Results Reference

3% alginate-RGD beadswith poroussurface.

Drops of alginate-RGD weredripped in into calciumchloride in glacial acetic acidsolution. Beads placed Undervacuum to create poroussurface.

Subcutaneousimplantationin rats.

Cellular migration of fibroblasts,endothelial cells, primarymacrophages and giant cells intobeads was shown. Little inflammationand good vascularization.

Loebsack etal. (2001)

1.75%alginateporous beads

Alginate solution was stirredintensively to incorporate air inthe solution. Drops of alginate-RGD were dripped in intocalcium chloride in glacialacetic acid solution. Beadsplaced Under vacuum to createporous surface.

Fibroblastcell cultureandsubcutaneousimplantationin rats.

Beads have an open pore structurewith a porosity of about 78%.Fibroblasts were found to evenlydistribute throughout the entire bead.Implantation showed that the porestructure remains intact and allowscell invasion. Good vascularizationwas observed.

Eiselt et al.(2000)

1-3% alginateporoussponges

Lyophilization with freezing at-18°C and in liquid nitrogen.

Fibroblastculture withinsponges.

Pore size depended on the alginatetype as well as the concentration ofalginate. Pore size of samples frozenin liquid nitrogen ranged from 60-200µm. The sponges with or without cellsin culture medium at 37°C retainedtheir original shape and configurationfor prolonged times.

Shapiro et al.(1996)

Porousalginate beadswith orwithouthyaluronicacid

Lyophilization with freezing at-20°C.

Chondrocyteculture withinsponges.

The pore size of the sponges was 174± 99 µm. Chondrocytes exhibited aspherical shape with a non-orientednetwork of actin filaments. Thepresence of hyaluronic acid did notchange the histological score.

Miralles et al.(2001)

Porousalginatesponges

Lyophilization with freezing at-18°C.

Hepatocyteculture withinsponges

The sponges had interconnected poreswith pore sizes ranging between 100-150 µm. Hepatocytes arranged insmall viable aggregates.


Porousalginatesamples

Lyophilization with freezing at-20°C, -30°C, and liquidnitrogen.

Hepatocyteculture.

They found no significant differencein pore size between the differentfreezing regimes. Hepatocytes showedspherical morphology and aggregatedinto clusters.

Zmora et al.(2002)

Table 3.4: An overview of studies concerning porous alginate scaffolds.

Loebsack et al. (2001) created a porous alginate scaffold, which was modified by aGGGGRGDY adhesion peptide. Alginate beads were prepared with the dripping methoddescribed above. The drops of alginate were dripped into a solution of 0.5M of calciumchloride in glacial acetic acid (9:1) and were placed under vacuum for four hours. This createdan alginate bead with a porous surface. These beads were implanted subcutaneously in rats.Histological evaluation of the samples after implantation revealed cellular migration offibroblasts, endothelial cells, primary macrophages, and giant cells into the interstices of thebeads. Little inflammation was observed and the samples were well vascularized. Eiselt et al.(2000) created more or less the same porous beads but included one extra step. Prior todripping the alginate drops into the calcium chloride glacial acetic acid solution, the alginatemixture was stirred extensively during 30 minutes to incorporate air into the solution. After thebeads were formed, they were placed under vacuum to remove the incorporated air bubblesin the beads. This led to an open interconnected pore structure with a porosity of about 78%.By seeding fibroblasts into the beads, they were able to demonstrate a uniform distributionand interconnection of the pores. Subcutaneous implantation revealed that the porousstructure remained intact and allowed cell invasion. Blood vessels containing erythrocyteswere observed in the polymer interior thus demonstrating that the pores were large enough toallow vascularization of the alginate beads. Shapiro et al. (1996) created porous alginatesponges with the freeze drying technique. The sponges were sterilized using an ethyleneoxide gas treatment. Through electron microscopy it was observed that the alginate spongesdisplayed a highly porous, well-interconnected pore structure. The pore size depended on thealginate type that was used as well as the concentration of alginate and the amount and typeof crosslinker and the freezing temperature. The average pore size of sponges frozen in liquidnitrogen ranged from 60 to 200 µm. The mechanical properties of the dry sponges were


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examined. The elastic modulus ranged from 120 to 380 kPa. Miralles et al. (2001) createdalginate sponges in the same way as Shapiro et al. (1996), but combined the alginate withhyaluronate. The pore size of these sponges was 174 ± 99 µm. Chondrocytes wereincorporated into the sponges and exhibited a spherical shape with a non-oriented and sparsenetwork of actin filaments. The presence of the hyaluronate did not influence the histologicalscore (cell density, cell morphology, proteoglycan deposition, collagen deposition, andnetwork organization) of the chondrocyte-sponge. Glicklis et al. (2000) also created alginatesponges in the same manner as the two previous authors and found a porous structure withinterconnected pores and pore sizes ranging between 100-150 µm in diameter. A culture ofhepatocytes was incorporated into the porous samples. The cells arranged in small viableaggregates. Zmora et al. (2002) examined the possibility of controlling the pore structure ofalginate sponges by using different freezing temperatures. Surprisingly enough, they didn’tfind a significant difference in pore size between the different freezing regimes. Aftercompression of the dry alginate sponges, they found a significant higher modulus of elasticity(1136 kPa) for samples frozen at -20°C than for samples frozen in liquid nitrogen (385 kPa).

3.2.5 Combining alginate with fibrin gel

Scaffold Method In vitro /in vivo

Results Reference

Fibrin beads Alginate and fibrin solutionwas dripped into calciumchloride solution containingthrombin. The alginate wasremoved after polymerizationwith sodium citrate.

Periosteal cellculture.

The number of cells increasedsignificantly compared to the initialvalues.

Spitzer et al.(2002)

Alginate,fibrin, andalginate-fibrinbeads

Alginate and fibrin solutionwas dripped into calciumchloride solution containingthrombin. The alginate wasremoved after polymerizationwith sodium citrate. Foralginate beads only alginatewas dripped into calciumsolution.

Chondrocyteculture

Significant higher cell proliferation inalginate-fibrin beads than in alginatebeads. In both cases the cell retained aspherical shape. In fibrin beads, thechondrocytes dedifferentiated intofibroblast-like cells.

Perka et al.(2000 and2001)

Alginatebeads infibrin gels

Alginate-cell solution wasdripped into calcium chloridesolution. After polymerization,the alginate beads wereencapsulated in fibrin gel.

Chondrocyteculture in thealginatebeads.

Cell proliferation was significantlymore pronounced in the cultureswhere alginate was surrounded byfibrin. Cells appeared in thesurrounding fibrin gel after two weeksof culturing. Collagen and aggrecanwere found in the fibrin after eightweeks.

Almqvist etal. (2001)

Table 3.5: An overview of studies concerning alginate-fibrin combinations.

Another way of enabling cell attachment to alginate is mixing the alginate with fibrin. Fibrindoes have the ability to bind cells. Spitzer et al. (2002) used alginate-fibrin beads to examineosteogenic differentiation of rabbit periosteal cells. Alginate and fibrinogen were mixed anddripped into a calcium chloride solution containing thrombin. After polymerization, the alginatewas removed from the beads by application of a solution of sodium citrate, sodium chloride,and EDTA, resulting in a porous fibrin bead with a diameter of 2.9 mm. Perka et al. (2001)used the same method to create fibrin beads and performed a cell study with chondrocytes onalginate beads, fibrin beads, and alginate-fibrin beads. They found a significantly higher cellproliferation in the composite alginate-fibrin beads than in alginate beads. In both types ofbead containing alginate, the cells retained a spherical shape throughout the culturing period.In the fibrin beads however, the chondrocytes dedifferentiated into fibroblast-like cells. Perkaet al. (2000) used the same three types of beads to examine chondrocyte proliferation andfound a significant higher rate of proliferation in fibrin beads as compared to alginate andalginate-fibrin beads. The morphology of the cells remained spherical in shape in the alginateand alginate-fibrin beads. In the fibrin beads however, the cells formed cytoplasmicextensions and dedifferentiated into fibroblast-like cells. Almqvist et al. (2001) examinedchondrocytes in alginate beads by dripping an alginate-chondrocyte solution into a calcium


35

chloride solution. After the calcium chloride was removed, the beads were washed withsodium chloride and cultured in well plates. Some of the beads were encapsulated in fibrin gelin order to follow the proliferation of chondrocytes in the beads and outgrowth of these cellsinto the surrounding fibrin. It was observed that the cells in the alginate beads proliferated to ahigh extent. Cell proliferation was significantly more pronounced in the cultures wherechondrocytes were cultured in alginate beads surrounded by fibrin gel. Cells appeared in thesurrounding fibrin gel after two weeks of culturing. Collagen and aggrecan were found in thesurrounding fibrin gel by week eight.


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Chapter 4

In this section the research concerning the tissue engineering of heart valves will bereviewed. First, the phenomenon of tissue engineering will be explained.

4.1 Tissue engineeringThe most common surgical procedure for end-stage valvular heart disease is valvereplacement, with more than 60,000 implantations in the United States and 170,000worldwide each year. In the United States, every year, approximately 15,000 infants are bornwith congenital cardiac malformations, many of which involve one or more cardiac valves[Fuchs, 2001; Langer, 1993]. To avoid the problems associated with artificial aortic valvereplacements described in chapter one, a new method called tissue engineering hasemerged. The replacements described earlier all consist of dead material and are unable toadapt to changing circumstances. Especially in young children the aortic valve replacementsneed to be flexible and able to adapt to the changing circumstances and must be able tosense and respond to the normal, biological signals for growth and development [Mann,2001]. In tissue engineering, the patient’s own cells are used so that there is no chance of animmune response. These autologous cells are seeded onto a biodegradable scaffold in theshape of a heart valve. Figure 4.1 schematically reviews the steps that are taken in the tissueengineering process.

Figure 4.1: The tissue engineering process.

The scaffold has to have a number of essential properties. It should be biocompatible,biodegradable, very porous, permeable to allow proper diffusion, and have the correct poresize for the candidate cells. The biodegradability of the scaffold allows it to be graduallyreplaced by new cells and extracellular matrix to form functional tissues. Furthermore, ascaffold should have the correct mechanical properties [Agrawal, 2001] and a suitable surface


37

chemistry, which influences the cell attachment, cell proliferation, cell migration, and celldifferentiation. A scaffold can be used to perform several additional functions, for instance, itcan act as a means to deliver an active protein, gene, or cell product via controllabledegradation, or as a means to orient or manipulate transplanted host tissues [Mol, 2001].

4.1.1 Research on the tissue engineering of heart valves

Fuchs et al. (2001) reviewed some of the work performed on the tissue engineering of heartvalves. They mention the first valve leaflet ever tissue engineered. The study used mixed cellpopulations of endothelial cells and myofibroblasts from ovine arteries. Polyglycolic acid(PGA) scaffolds were seeded with myofibroblasts and then were subsequently seeded withendothelial cells. The constructs were implanted in place of the native right posterior leaflet ofthe pulmonary valve in sheep. No stenosis was observed and only moderate regurgitation.Other studies, using both cell-seeded scaffolds and only the scaffold material (PGA) revealedthat the polymer implants were completely degraded by eight weeks, whereas the tissueengineered leaflet persisted. They found elastin and a progressively increasing collagencontent in the tissue engineered leaflets. Further work concentrated on the development of atrileaflet valve that could be tested in a pulsatile flow bioreactor. The PGA scaffold provedunsuitable for the tissue engineering of a trileaflet construct because of the material’sstiffness. Other scaffold materials like poly-4-hydroxybutyrate (P4HB) demonstrated aconsiderable amount of cell attachment and considerable amount of collagen produced by thecells was found.

Sodian and Sperling (2000) investigated the possibility of thermoplastic biocompatiblepolyesters known as polyhydroxyalkanoates (PHAs) to be used as scaffolds for the tissueengineering of a trileaflet heart valve. They used a salt leaching technique to create a porousthree-dimensional structure from the PHA. They created pores that varied in size between 80and 200 microns. Cells were able to attach to the polymer and formed a confluent layer afterincubation and they concluded that PHAs could be used to fabricate a three-dimensional,biodegradable heart valve scaffold. Sodian and Hoerstrup (2000) evaluated three differentmaterials as a scaffold for tissue engineering of heart valves. They tested polyglycolic acid(PGA), polyhydroxyalkanoate (PHA), and poly-4-hydroxybutyrate (P4HB). This is quiteconfusing, because both PGA and P4HB belong to the family of polyhydroxyalkanoates(PHAs). They modified the PHA and P4HB scaffolds with a salt leaching technique to create aporous matrix. The constructs were seeded with ovine vascular cells for eight days and thenexposed to continuous flow for one hour. They analyzed the samples by comparing the cellproliferation, cell attachment, and collagen production before and after the flow exposure.They found that there were significant more cells on PGA compared with PHA and P4HBbefore flow exposure, although after flow exposure there were no differences found betweenthe three scaffolds in cell number. However, cell attachment and collagen production weresignificantly higher on PGA samples compared to PHA and P4HB. PHA and P4HB alsodemonstrated a considerable amount of cell attachment and collagen development and sharethe major advantage that both materials are thermoplastic, making it possible to mold theminto the shape of a functional scaffold for the tissue engineering of heart valves. Hoerstrupand Zund (1999) tested two methods to stimulate the production of extracellular matrix ofhuman myofibroblasts on control scaffolds. To stimulate collagen production, one series wascultured with L-ascorbic acid 2-phosphate. In a second series, the seeded scaffolds weresubjected to tension by mounting them on a frame. Collagen content of the framed scaffoldswas ten times higher than that of the control group and six times higher than in the unframedscaffolds grown with ascorbic acid, indicating the importance of mechanical stimulation of aseeded tissue engineering scaffold to stimulate the production of extracellular matrix. In orderto address the problems of the inadequate mechanical properties of tissue engineered heartvalves, Hoerstrup and Sodian (2000) developed an in vitro cell culture system that providesphysiological pressure and flow of nutrient medium to the developing valve constructs. A newdynamic bioreactor (figure 4.2) allows adjustable pulsatile flow and varying levels of pressure.Thus, this bioreactor delivers the mechanical stimulation the tissue engineering constructsneed to produce extracellular matrix.


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Figure 4.2: A bioreactor with a pulsatile flow chamber. Adapted from Hoerstrup et al. (2000).

Sodian and Hoerstrup (2000) constructed a three-dimensional biodegradable andbiocompatible trileaflet heart valve scaffold from a porous PHA. The scaffold consisted of acylindrical stent and three leaflets which were 0.3 mm thick. The leaflets were attached to thestent by thermal processing techniques. The porous scaffold (pore size 100 to 240 microns)was seeded with vascular cells and placed into a pulsatile bioreactor for 1, 4, and 8 days. Itwas shown that the cells grew into the pores and formed a confluent layer after incubationand pulsatile flow exposure. The cells were mostly viable and formed connective tissuebetween the inside and outside of the porous heart valve scaffold. Cell proliferation and thecellular capacity to generate collagen was demonstrated through DNA and hydroxyprolineassay respectively.

Another study performed by Sodian and Hoerstrup (2000) was designed to evaluate theimplantation of a whole trileaflet tissue engineered heart valve in the pulmonary position in alamb model. They used PHA with pore sizes from 180 to 240 microns and seeded vascularcells on the heart valve scaffold. The constructs were incubated for 10-12 days in cell culturemedium supplemented with L-proline, L-alanine, glycine, and L-ascorbic acid. After the 10-12days the constructs were additionally seeded with venous endothelial cells and left toincubate for one day. No preconditioning in a pulsatile bioreactor was performed. Theconstructs remained in the animals up to seventeen weeks. After the implantation time thevalves were harvested from the animals. The valves showed minimal regurgitation and thetissue engineered constructs were covered with tissue, and there was no thrombus formation.Scanning electron microscopy showed laminated fibrous tissue with predominantglycosaminoglycans as extracellular matrix. Hydroxyproline assays demonstrated a collagenincrease of about 70% from week one to week seventeen. DNA assays showed a comparablenumber of cells in all explanted samples.

Hoerstrup and Sodian (2000) again demonstrated the production of a three-dimensionalscaffold in the form of a heart valve, but this time used a combination of PGA and P4HB tothermoplastically mold the mixture in the form of a trileaflet valve. The construct was seededwith myofibroblasts and endothelial cells and matured for fourteen days in a pulsatile flowbioreactor under gradually increasing flow and pressure conditions. Figure 4.3 shows thetissue engineered heart valve after fourteen days in a pulsatile flow bioreactor.


39

Figure 4.3: The tissue engineered heart valve after fourteen days of pulsatile flow in a bioreactor.Adapted form Hoerstrup et al. (2000).

In the bioreactor, the valve showed synchronous opening and closing of the leaflets under lowand high pressure conditions. The valves stayed in the bioreactor for up to 28 days. Thetissue engineered leaflets of the valve showed cellular tissue organized in a layered fashionwith a dense outer layer and less cellular activity in the deeper portions after fourteen days inthe bioreactor. Collagen content was 129% compared to native valves (100%) at fourteendays and leveled of to 85% at four weeks. The static controls showed less tissue formationand organization at all time points. ESEM demonstrated dense tissue and a confluent smoothsurface with cell orientation in the direction of flow after seven days, whereas the staticcontrols showed a rough surface at all time points. When the results of collagen content whena bioreactor was used (129%/85%) were compared to the results from static culturing bySodian and Hoerstrup (2000) described above (70%), it can be seen that mechanicalstimulation in a bioreactor does have a positive influence on the collagen production.The constructs were then implanted in lambs at the position of the pulmonary valve. Theimplanted constructs were observed with echocardiography and demonstrated mobilefunctioning leaflets without evidence of thrombus, stenosis, or aneurysm formation up totwenty weeks of implantation. Histology showed a uniform laminated structure withprogressive thinning and organization of the cuspal structure in time. The leaflets werelayered with a loose spongy layer mainly consisting of glycosaminoglycans on the ventricularside and a fibrous layer containing collagen on the arterial side. Collagen content was 140%after four weeks and reached a maximum at 180% after eight weeks. The tensile strength ofall implanted tissue engineered valve leaflets was initially higher than that of native tissue anddecreased over the follow-up period to be comparable to native values.Schnell and Hoerstrup (2001) explored the optimal cell source for cardiovascular tissueengineering and found that easy-to-access venous myofibroblasts showed excellent in vitrotissue generation. Collagen formation and mechanical properties were superior to aortictissue derived constructs.A recent publication by Hoerstrup et al. (2002) used a trileaflet heart valve scaffold made fromPGA/P4HB composite material. They used human marrow stromal cells to seed onto thescaffolds and cultured the constructs for fourteen days in a pulsatile bioreactor with graduallyincreasing nutrient media flow and pressure conditions. The constructs showed mechanicalproperties and morphological features resembling native heart valves and it was concludedthat human marrow stromal cells represent a promising cell source for cardiovascular tissueengineering purposes.


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Discussion and future research

From this review report, it becomes clear that the tissue engineering of aortic heart valves is amulti-disciplinary effort. It is essential to have knowledge of the natural valve: The mechanicalproperties of natural aortic valve tissue and the anatomical and geometrical characteristics(described in chapter one) are important, a tissue engineered construct must resemble thenatural aortic valve in these properties. In the tissue engineering approach, a scaffold materialis used to support the cells in the first stadium of development and eventually the scaffoldmaterial will degrade and the, by the cells, produced matrix will have taken over themechanical load. It is essential for cells to be attached to the scaffold material; withoutattachment the genesis and maintenance of both three-dimensional structure and normalfunction in tissues would be impossible. Cell attachment is critical for the cell to have controlover its cycle, control over programmed cell death (apoptosis), control over shape, growth,survival, and gene expression. Cells “communicate” with the extracellular matrix through focalcontacts. At focal contacts, clusters of specific adhesion receptors called integrins bindexternally to extracellular matrix molecules and internally to cytoplasmic proteins that in turnbind to the cytoskeleton. Integrin signal transduction relies on the activation of tyrosinekinases. Upon activation of these kinases, the focal adhesion kinase (FAK) pathway or Shcpathway is triggered. Both pathways lead to the activation of the mitogen-activated proteinkinase (MAPK) cascade. In this cascade, proteins are activated upon phosphorylation andcan subsequently move into the nucleus of the cell and act as a gene regulatory protein. Thisprocess is described in chapter two of this report.In this research an algae derived material called alginate is used in the production of ascaffold for the tissue engineering of an aortic heart valve. Regretfully, alginate misses thecapacity to interact with cells. A way to overcome this problem is to covalently modify thealginate molecules with an Arginine-Glycine-Aspartic acid (RGD) sequence containingpeptide. RGD is a specific amino-acid sequence that is recognized by several members of theintegrin family. The work of Rowley et al. (1999, 2002) revealed that modification of alginatewith an RGD-containing peptide leads to the binding of cell to the scaffold. On the other hand,Mann et al. (1999, 2001, and 2002) demonstrated that a too strong attachment of the cell tothe scaffold material leads to a decrease in matrix production. This can however becompensated by the addition of specific growth factors tethered to the scaffold or dissolved inthe medium.

Several studies were performed in which an actual heart valve was tissue engineered(Hoerstrup, Sodian, Turina, Mayer, 1999-2002). These aortic heart valves were first culturedin vitro in a pulsatile bioreactor. The mechanical stimulus in such a bioreactor triggers theproduction of extracellular matrix by the cultured cells. Although the morphologicalappearance of the tissue engineered valve looks promising, the weakness of the tissue is themain shortcoming of these constructs. The valves are only suited for implantation at theposition of the pulmonary valve and not as replacement for the aortic valve, because they cannot withstand the magnitude of the mechanical forces present there.

The work and goal of this graduation research is part of the investigations to accomplish astronger tissue engineered aortic valve construct. This can possibly be achieved by mixingdifferent types of scaffold material. For instance, a gel bulk structure combined with a networkof polymer fibers. In this research one of the possible candidates for the gel component isalginate. From the literature described in chapter three it becomes clear that alginate is easyto manipulate and problems associated with it can easily be solved. Therefore the goal of thisresearch is to create a suitable alginate scaffold construct, which in the future could act as aheart valve scaffold or as part of a heart valve scaffold. It must live up to certain requirements.It must have the correct mechanical properties and have a highly interconnected porousstructure. Furthermore, it must have interaction with cells. A method of creating porosity is


41

lyophilization. Lyophilization comprises two steps; the first is freezing of the sample. Thesecond is the sublimation of the ice crystals. Applying different freezing temperatures, asexplained by Kang et al. (1999) can control the pore size. Furthermore, the pore orientationcan be influenced by application of a temperature gradient over the sample during thefreezing. In this research project, it is hypothesized that as with gelatin (examined by Kang etal. (1999)), the pore size in the alginate constructs will decrease with decreasing freezingtemperature. The pore structure and size will be evaluated by light microscopy, micro-CTimaging (computed tomography), and environmental scanning electron microscopy.Mechanical properties of alginate samples will be tested on a compression/tension bench.Two different properties will be tested. The strength of the samples will be measured andrepresented by the modulus of elasticity (or Young’s modulus). Furthermore, the maximumstrain of samples of different concentrations of alginate will be measured. Special attentionwill be given to freeze dried samples, because it will be important to know to what extent thecreation of a porous structure has influence on the mechanical properties of the alginatesamples.Alginate molecules will be modified with RGD-containing peptides in order to overcome theproblem of the lack of cellular interaction of the scaffold material. Aqueous carbodiimidechemistry will be applied to covalently link the RGD-containing peptide to the alginatemolecules. A combination between fibrin gel and alginate can be another solution for theproblem of the lack of cellular interaction and will be worked out in this research. Cell studieswill be performed on non-porous, unmodified alginate surfaces to confirm the lack of cellularinteraction of the alginate molecules. Solvent swelling experiments with cells will beperformed on freeze dried, unmodified alginate samples. By dropping the cell-mediumsolution on the dry samples, which will soak up the cell-medium solution, solvent swelling willbe performed. Solvent swelling experiments will also be performed with (G)RGD-modified,freeze dried alginate samples and freeze dried fibrin-alginate samples.


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