Tissue Engineering: Seeding Cells with Synthetic Materials for Regeneration, Notas de estudo de Engenharia de Produção

Baixe Tissue Engineering: Seeding Cells with Synthetic Materials for Regeneration e outras Notas de estudo em PDF para Engenharia de Produção, somente na Docsity! nature materials | VOL 8 | JUNE 2009 | www.nature.com/naturematerials 457 review article Published online: 21 may 2009 | doi: 10.1038/nmat2441 A human embryo in its first eight weeks of life undergoes an extraordinary transformation from a single cell to a 3-cm-long fetus with a beating heart, gut, nervous system, and limbs with fingers and toes. This progression involves massive growth, physical folds and twists, and myriad cellular and molecular events of breathtaking complexity; yet it is the ultimate goal of tissue engineering (TE) to recreate some of these processes in microcosm, to replace and regenerate lost tissue. At last the field has entered a period of fruition, and seems set to realize its potential to treat a multitude of debilitating and deadly conditions such as myocardial infarction, spinal injury, osteoarthritis, osteoporosis, diabetes, liver cirrhosis and retinopathy. The general strategy is usually to seed cells within a scaffold, a structural device that defines the geometry of the replacement tissue and provides environ mental cues that promote tissue regeneration. TE skin equivalents have been in clinical use since 1997 (ref. 1) and a fast-growing arsenal of replacement devices is in clinical trials or already approved as therapies for tissues includ- ing cartilage, bone, blood vessel and pancreas (Table 1). In two recent high-profile studies, seven patients benefited from TE blad- ders2, and a 30-year-old woman became the first person to receive a TE tracheal segment, a procedure that saved her left lung3. Aside from the obvious human benefits, tissue engineering could bring substantial financial rewards to those who succeed in trans- lating this new technology to the clinic. Sales of regenerative bio- materials already exceed US$240 million per annum4 and the wider markets that tissue engineering taps into are colossal: costs related to organ replacement account for 8% of global healthcare spending, and by 2040 as much as 25% of the US GDP is expected to be related to healthcare5. Nevertheless, if the short history of industrial tissue engineering has taught us anything, it is that the provision of effec- tive products is not in itself sufficient to ensure commercial success (Fig. 1). Early TE efforts were plagued by product issues related to scale-up, shelf-life, quality control and distribution, and suffered from inappropriate business models and withdrawal of private finance in the early 2000s1,6. Since then the field has matured, evi- denced by the return of large-scale investment and the first regen- erative medicine companies becoming profitable4. Alongside these positive developments, progress in biomaterials design and engineering are converging to enable a new generation of instructive materials to emerge as candidates for regenerative medi- cine. Which of these materials compete successfully in the market will depend on a combination of clinical performance, marketing and cost-effectiveness. A central dilemma is that to influence cell behaviour, scaffolding materials must bear complex information, Complexity in biomaterials for tissue engineering elsie s. Place1,2, nicholas d. evans1,2 and molly m. stevens1,2 The molecular and physical information coded within the extracellular milieu is informing the development of a new generation of biomaterials for tissue engineering. Several powerful extracellular influences have already found their way into cell- instructive scaffolds, while others remain largely unexplored. Yet for commercial success tissue engineering products must be not only efficacious but also cost-effective, introducing a potential dichotomy between the need for sophistication and ease of production. This is spurring interest in recreating extracellular influences in simplified forms, from the reduction of biopolymers into short functional domains, to the use of basic chemistries to manipulate cell fate. In the future these exciting developments are likely to help reconcile the clinical and commercial pressures on tissue engineering. coded in their physical and chemical structures. On the other hand, financial considerations dictate that complexity must be kept to a minimum. Clearly there is a danger, by over-engineering devices, of making their translation to clinical use unlikely. The solutions to this challenge lie at every phase of product development, begin- ning with identifying the simplest functional performance required to resolve a defined clinical problem. The ambitious early aims of reconstructing entire organs have largely given way to smaller, more attainable goals: for example, rather than trying to replace an entire heart, clinical advances in cardiac repair focus on TE coro- nary arteries, valves and myocardium. Organogenesis Inc. and Advanced Tissue Sciences Inc. suffered heavily as a result of their overestimating the number of chronic wounds cases that were best solved by high-tech, TE skin substitutes (respectively, Apligraf and Dermagraft; Dermagraft is now produced by Advanced Biohealing) as opposed to acellular products that aid ongoing repair6 (Table 1, Fig. 1). Similarly, an emerging philosophy in tissue engineering is that rather than attempting to recreate the complexity of living tissues ex vivo, we should aim to develop synthetic materials that establish key interactions with cells in ways that unlock the body’s innate powers of organization and self-repair. In this review we will consider how this can be achieved, emphasizing how even relatively simple engineering solutions can deliver considerable functional benefits. Along the way we will explore how some of these princi- ples have been applied to specific scientific and commercial tissue- engineering challenges. regenerative potential of tissues Even without any therapeutic intervention, living tissues can have a staggering capacity for regeneration. For example, the human liver will regrow to its original size even when more than 50% of its mass is excised7. This has been taken to the extreme in rats, where one group has reported that a single rat’s liver was able to regenerate fully following each of 12 sequential hepatectomies, a finding that can be explained by the high replicative potential of the cell types that make up the liver. Several other tissues — bone and skin, for example — also have an innate capacity to regener- ate to fill injuries below a critical size, helped by local or recruited stem cells. The clinical potential of stem cells has long been rec- ognized by haemato logists, who in the 1960s showed that trans- planted haemato poietic (literally ‘blood-making’) stem cells from the bone marrow of a healthy mouse could replace the destroyed immune system of another mouse, paving the way for a cure for leukaemia8,9. The discovery of other types of cell with multilineage 1Department of Materials; 2Institute for Biomedical Engineering, Imperial College London, London SW7 2AZ, UK. e‑mail: m.stevens@imperial.ac.uk nmat_2441_JUN09.indd 457 12/5/09 15:34:03 © 2009 Macmillan Publishers Limited. All rights reserved 458 nature materials | VOL 8 | JUNE 2009 | www.nature.com/naturematerials review article NaTure maTerIalS doi: 10.1038/nmat2441 potential has since followed, including neural stem cells from the brain, and mesenchymal stem cells, which can differentiate into bone, fat, cartilage and muscle cells10,11. Indeed, more recent evidence suggests that stem cells or progenitor cells can be isolated from almost every tissue of the body12,13. Under the correct condi- tions, these cells can be stimulated to form new tissue, as we recently demonstrated using a simple biomaterials-based approach. Here, either calcium-crosslinked alginate gels or modified hyaluronic acid gels were injected into an artificial space between the tibia and the periosteum, the fibrous outer lining of bone. This stimu- lated bone and cartilage formation from resident progenitor cells in the inner layer of the periosteum14, illustrating that complex tissues can be generated from relatively simple materials by using the body as a ‘bioreactor’. Table 1 | Commercial tissue engineering products and biomaterials at various stages of development. tissue Product regulatory status description material Cells use Form Sy nt he tic re so rb ab le a ni m al d er iv ed Pl an t o r b ac te ria de riv ed H um an d er iv ed G ro w th fa ct or a llo ge ni c a ut ol og ou s Skin TransCyte, Advanced Biohealing 1997 Nylon mesh coated with porcine collagen, containing non‑viable human fibroblasts, with upper layer of silicon ✓ ✓ ✓ ✓ Burns Sheet Apligraf, Organogenesis 1998 Lower layer of human fibroblasts and bovine collagen, upper layer of keratinocytes ✓ ✓ ✓ Leg ulcers Sheet Dermagraft, Advanced BioHealing 2001 Cryopreserved human fibroblasts on a polyglactin 910 (2‑hydroxy‑propanoic acid polymer with polymerized hydroxyacetic acid) mesh ✓ ✓ ✓ Diabetic foot ulcers Sheet ICX‑SKN, Intercytex Phase II Allogenic fibroblasts and human collagen with additional layer of keratinocytes ✓ ✓ ✓ Burns and acute wounds Sheet Integra Dermal Regeneration Template, Integra Lifesciences 1996 Porous bovine collagen crosslinked with chondroitin‑6‑sulphate with upper layer of silicon ✓ ✓ ✓ Burns Sheet Integra Flowable Wound Matrix, Integra Lifesciences 2007 Granulated bovine collagen crosslinked with chondroitin‑6‑sulphate ✓ ✓ Ulcers Gel Oasis Wound Matrix, Healthpoint 2006a Decellularized porcine small intestinal submucosa ✓ ✓ Burns, ulcers, other wounds Sheet PriMatrix, TEI Biosciences 2008 Decellularized fetal bovine skin ✓ ✓ Wounds Sheet Xelma, Molnlycke 2005 EU ECM protein (amelogenins) in propylene glycol alginate carrier ✓ ✓ ✓ Leg ulcers Gel Bone INFUSE Bone Graft, Medtronic 2002 Bovine type I collagen sponges soaked in rhBMP‑2 in LT‑CAGE Lumbar Tapered Fusion Device ✓ ✓ ✓ ✓ Spinal fusion Solid OP‑1, Stryker 2001 Bovine type I collagen with rhBMP‑7 ✓ ✓ ✓ Bone injury Paste PuraMatrix, 3DM Preclinic Synthetic 16‑amino‑acid peptide, forming nanofibres ✓ ✓ Dental bone defects Gel Vitoss Scaffold FOAM, Orthovita 2004 Porous foam comprising β‑TCP and bovine type I collagen ✓ ✓ ✓ Bone injury Foam Bioset IC, Pioneer surgical 2008 Human demineralized bone matrix with bovine bone chips in type I collagen carrier ✓ ✓ ✓ Bone injury Paste FortrOss, Pioneer Surgical 2008 Nanocrystalline hydroxyapatite and E‑matrix (porcine collagen co‑polymerized with dextran) ✓ ✓ ✓ ✓ Bone injury Paste Regenafil, Regeneration Technologies/Exatech 2005 Human mineralized bone matrix in porcine gelatin carrier ✓ ✓ ✓ Bone injury Paste GEM 21S, BioMimetic Therapeutics 2005 β‑TCP particles and recombinant human platelet‑derived growth factor‑BB (PDGF‑BB) ✓ ✓ ✓ Dental bone/ gum defects Paste BCT001, Bioceramic Therapeutics Preclinic Strontium releasing bioactive glasses ✓ ✓ Bone defects Granules, paste nmat_2441_JUN09.indd 458 12/5/09 15:34:03 © 2009 Macmillan Publishers Limited. All rights reserved nature materials | VOL 8 | JUNE 2009 | www.nature.com/naturematerials 461 review articleNaTure maTerIalS doi: 10.1038/nmat2441 mater layer of the brain meninges following a craniotomy (Table 1). In a further development, last year’s transplanted TE airway con- firmed this approach as being at the forefront of whole-organ tissue engineering 3,31. The scaffold in this case was a decellularized human donor trachea that was repopulated with the patient’s own cells expanded from biopsy. In contrast with traditional transplant surgery, the decellularization protocol solved the problem of tissue rejection by removing virtually all traces of human leukocyte antigens (the proteins that to a large extent determine tissue compatibility), with the consequence that the patient required no immuno suppressive drugs. As well as immediately restoring airway patency, the device facilitated the rapid development of an internal cellular lining and blood vessel network. Although we focus here on scaffolds designed and assembled in ‘bottom-up’ mode in the laboratory, it is apparent that both lab-built scaffolds and decellularized tissues offer distinct and important benefits for tissue engineering, and equally, that nei- ther approach represents a universal biomaterials solution. substituting physical aspects of the extracellular matrix Typically, biomaterials-engineering approaches focus on a few mechanisms (chemical or physical) by which ECM influences cells, and attempt to present these influences effectively for a given tissue. Regardless of the chemistry that we apply within scaffolds, the con- struct must usually also provide some level of physical support from the moment of implantation, to assist tissue function while new matrix is being deposited32–35. For example, the extreme softness of the lamina propria of the human vocal fold (elastic modulus E = 100–1,000 Pa) is essential for proper phonation, and its function is easily impaired by scarring. This has prompted the development of soft (E ≈ 500 Pa), highly elastic gels of double-crosslinked hyaluronic acid microparticles for vocal-fold tissue engineering36. The par- ticles are synthesized by crosslinking with divinyl sulphone, then surface-oxidized and lightly crosslinked together using hyaluronic acid modified with adipic dihydrazide. The gels have controllable viscoelasticity, and a reduction in dynamic viscosity with frequency occurs at a similar rate to that of human vocal-fold mucosa. In many cases, physical demands on scaffolding materials are complicated by the anisotropy inherent in most living tissues (con- sider the parallel arrangement of collagen fibres within tendons and the concentrically layered sheets of the intervertebral disk). Nevertheless, engineering solutions need not be costly or comp- licated: substituting a rotating for a static collector yields orien- tated electrospun fibres37, and crosslinking hydrogels under strain can result in highly biomimetic anisotropic mechanical properties. For instance, thermal cycling of poly(vinyl alcohol) leads to the growth of crystallites that function as physical crosslinks, leading to gelation. Early in the crosslinking process, these crystallites can be aligned by applying strain, the degree of which dictates the level of anisotropy. Thermal cycling is recommenced, with the number of cycles determining the overall amount of crosslinking and thus stiffness. By optimizing these two parameters, hydrogels with aniso- tropic stiffnesses closely resembling those of porcine aorta have been developed38. Tissues are also heterogeneously organized into mechanically distinct zones, for example the superficial, transitional, Achieving a strong bond between mechanically dissimilar materials is as much a challenge in tissue engineering as in other branches of engineering. The morphological specialities and mechanical gra- dients seen at interfaces between musculoskeletal tissues in vivo reduce impedance mismatch and minimize stress concentrations as loads are redistributed, yet even with nature’s elegant solutions, many chronic musculoskeletal injuries occur at tissue boundaries. Unsurprisingly, rupture at insertion sites is also the most common cause of failure of ligament and tendon grafts131. Although aware- ness of this problem is growing, most orthopaedic TE devices do not feature distinct transition zones to improve load transfer between tissues. This includes most osteochondral plugs — bilami- nar bone and cartilage TE constructs that have been developed to improve the assimilation of cartilage into joints. Here, the accu- mulation of matrix can effect good integration between the two phases132,133, but few examples contain regions of calcified cartilage reminiscent of the ‘tidemark’ seen adjacent to subchondral bone in vivo. An interesting exception is an osteochondral graft consist- ing of a ‘bone’ component of hydroxyapatite populated with BMP-7 transduced fibroblasts (connective tissue cells), and a poly(lactic acid) sponge seeded with chondrocytes (cartilage cells). Pockets of mineralized cartilage were seen at the boundary between the two layers of this scaffold134. Conversely, trilaminar scaffolds by design possess a middle layer with a distinct composition135 and/ or seeded with different cell types. Any combination of cells can be straight forwardly zoned within hydrogels at the point of fabrica- tion by the layer-by-layer partial photo-polymerization of cell and macro molecular precursor suspensions136. Constructs resembling ligament insertion sites, wherein bone and ligament are united by means of fibrocartilage (Fig. B1), have been produced by seeding fibroblasts, chondrocytes and osteoblasts (bone cells) separately into the three layers of a preformed scaffold137. Another strategy uses just one cell type, namely fibroblasts, to produce scaffolds with an internal gradient of mineralization. Retrovirus encoding the bone-specific transcription factor Runx2 was immobilized on a layer of poly(l-lysine). The thickness of the poly(l-lysine) layer could be graduated by dip-coating collagen scaffolds, leading in turn to a tapering of retro virus concentration, osteoblastic gene expression, mineralization and stiffness138. Although the ligament components in these examples were not optimized for immediate tensile load bearing, TE ligaments with high tensile toughness (such as braided polymers139) have been developed34. It will be interesting to observe how, in the future, these two strands of ligament tissue engineering will be united. Box 1 | Challenges in interface engineering for orthopaedics. F BV Fb Cc Ob a CF B L T b T L F CF B Figure B1 | Histology of interface between bone and ligament. a, Schematic; b, photomicrograph of tibial insertion of rabbit anterior cruciate ligament. Adapted with permission from ref. 144. © 1996 Wiley. Ligament (L) insertions occur by means of fibrocartilage, which is divided at the tidemark (T, black line in b) into non‑calcified (F) and calcified (CF) regions. The calcified fibrocartilage interdigitates with the underlying subchondral bone (B). Fb, fibroblast; Cc, chondrocyte; Ob, osteoblast; BV, blood vessel. nmat_2441_JUN09.indd 461 12/5/09 15:34:17 © 2009 Macmillan Publishers Limited. All rights reserved 462 nature materials | VOL 8 | JUNE 2009 | www.nature.com/naturematerials review article NaTure maTerIalS doi: 10.1038/nmat2441 radial and tight zones of cartilage, some implications of which are discussed in Box 1. Beyond their structural and biomechanical roles, physical prop- erties also influence many aspects of cell behaviour. In one recent study, scaffold geometry was used to align cardiac muscle cells, to elicit directional contractions that are essential for efficient blood transfer39. The crosslinking of microstructured honeycomb poly(glycerol sebacate) sheets was optimized to mimic the aniso- tropic stiffness of rat ventricular myocardium. These sheets pos- sessed microscale pores in the form of two overlapping squares tilted at 45°, which directed the alignment of neonatal rat heart cells. The resulting constructs displayed anisotropic electrical exci- tation thresholds as a result of this long-range order. It is now well known that scaffolds that provide a particular physical environment can be cell-instructive as well as (potentially) contributing towards the physical function of the tissue. The stiffness of ECM alone can have effects at a transcriptional level, determining whether stem cells make the decision to turn into cells as functionally diverse as nerve and bone tissue40,41, and the importance of other physical features such as topography and three-dimensionality of the matrix has been reviewed or demonstrated elsewhere42–44. These physical factors continue to be investigated intensively and provide a com- plementary approach to the provision of molecular information to cells inside scaffolds. engineering bioactive scaffolds Tissue-engineering scaffolds can be designed to interact with cells by emulating key molecular features of the ECM. ECM contains many macromolecules such as proteoglycans, collagens, laminins, fibronectin and sequestered growth factors, and it is primarily this molecular information that confers its bioactivity. For example, the sequences of many ECM proteins are recognized by dimeric cell-surface receptors known as integrins. Binding of integrins to ECM molecules can trigger a cascade of signalling events lead- ing ultimately to gene expression. The gallery of ECM proteins presented to cells in a given tissue is likely to be critical in deter- mining how cells behave within that tissue. As integrins are dim- ers of alpha and beta subunits, they can associate in a variety of combinations, and thus bind to a diverse range of ligands. The indispensable nature of integrins and the ECM is demonstrated by genetic knock-outs: the absence of the integrin β1 subunit45 or of fibronectin46 is lethal at the early embryonic stage. One way to pro- vide sites for integrin attachment in scaffolds is to include purified ECM proteins such as collagen or fibrin. Table 1 demonstrates the success of this strategy, particularly regarding the use of collagen (and gelatin, a low-cost denaturation product of collagen), which is present in commercial products for most of those tissues where TE replacements are available. Another widely incorporated ECM molecule is hyaluronic acid, a polysaccharide that is deposited at sites of wound repair and also during morphogenesis—precisely the processes that we wish in some way to emulate. Hyaluronic acid is used therapeutically as a putative lubricating factor in arthritic joints (for example Synvisc; Table 1), although here and in tissue engineering it may also interact directly with cells through CD44 and other cell-surface receptors47. The ubiquity of these ECM molecules has prompted the development of Extracel (Table 1)48, a modular ECM system based on derivatives of gela- tin and hyaluronic acid. The components are combined in varying proportions to suit many different cell-culture and in vivo appli- cations, crosslinked with poly(ethylene glycol) (PEG) diacrylate and supplemented with additives such as heparin-bound growth factors as needed. For example, whereas equal quantities (weight for weight) of gelatin and hyaluronic acid are suitable for most tis- sue applications, a specialized composition enriched in hyaluronic acid (95:5 w/w hyaluronic acid/gelatin) is required for vocal-fold repair. This adaptability is especially valuable for niche markets, where the costs of development and manufacturing can be high relative to volume of sales. Matrix constituents have disadvantages associated with purifica- tion and processing, and coupled with the desire for greater control over material properties, this has led to the investigation of fully synthetic bioactive systems. The ability to functionalize bioinert sub- stances could improve the suitability for tissue engineering of a host of materials with remarkable properties, such as high-strength synthetic polymers and (nano)composites for bone tissue engineering 49,50. Of particular interest are systems amenable to minimally invasive deliv- ery, including injectable or shape-memory materials that gel or regain their original form in response to stimuli such as ultraviolet illumi- nation or physiological conditions (temperature, pH or solvent)51–54. Materials that do not adsorb protein, such as PEG gels, can effectively be used as a blank template on which to confer bioactivity with the minimum amount of modification. In some cases specific functions of biopolymers can be attributed to small functional domains which may be included in place of the full protein55,56. The best known of these is the integrin-binding arginine–glycine–aspartic acid (RGD) sequence found in many ECM proteins, including fibronectin, laminin, collagen type IV, tenascin and thrombospondin57–59. RGD modification alone is sufficient to transform alginate from a relatively inert polysaccharide into a substance that supports the formation of convincing growth-plate-like structures when mouse osteoblasts and chondrocytes are co-cultured within it60. Adhesive and informational peptides found in ECM are reviewed elsewhere61,62. In addition to providing cell-directing elements, ECM is itself highly responsive to the actions of cells. It is frequently remodelled by cells during development, homeostasis and healing, a process that involves digestion by a variety of proteases (such as cathep- sins and matrix metalloproteases) followed by deposition of fresh matrix. Many scaffolds undergo a hydrolysis route to degradation which can lead to sudden loss of mechanical strength and struc- ture. Cell-mediated scaffold degradation is more likely to generate a material temporal profile in tune with the generation of new tissue. In this approach, pioneered by the Hubbell group, the materials are crosslinked by enzyme-degradable peptide sequences, and a combi- nation of cell-mediated degradation and integrin binding allows the cells to migrate through the gel in a process reminiscent of tissue remodelling63. Cleavage sequences can also be incorporated into multidomain peptides. One example is a recombinant, crosslink- able elastin-like protein which harbours an adhesion motif (REDV) and an elastase-sensitive sequence. Cleavage of the latter yields a bioactive Val-Gly-Val-Ala-Pro-Gly (VGVAPG) fragment intended to stimulate cell proliferation and improve tissue repair64. This func- tionality mimics the multilayered bioactivity of the ECM, whereby enzymatic remodelling can liberate ‘cryptic sites’ contained within the amino acid sequences of ECM proteins. In fact it is becoming increasingly clear that fragments of many ECM proteins possess bioactivity — with effects ranging from cell migration to differentia- tion, proliferation and angiogenesis — which only becomes recognizable to the cells when the ECM is modified in some way65. Physical manipulation by cells can exert effects on the ECM: the interaction of integrins with fibronectin induce conformational changes of this large molecule, exposing cryptic sites which allow it to self-assemble66,67. In another example, myofibroblasts, cells implicated in wound healing, can release sequestered transform- ing growth factor beta (TGF-β) from its latent binding protein by pulling on the ECM, a process thought to be integrin-dependent68. Once relinquished, TGF-β is a powerful regulator of cell function, even to the extent of determining whether a cell lives or dies. soluble signalling molecules Growth factors, such as TGF-β and BMPs, are important signalling molecules in both tissue healing and development. Some of these proteins act as ‘morphogens’, determining the spatial arrangement nmat_2441_JUN09.indd 462 12/5/09 15:34:17 © 2009 Macmillan Publishers Limited. All rights reserved nature materials | VOL 8 | JUNE 2009 | www.nature.com/naturematerials 463 review articleNaTure maTerIalS doi: 10.1038/nmat2441 of cell types within the developing embryo. They may also have profound effects in the control of tissue regeneration. For example, injured muscle tissue secretes several Wnt proteins, stimulating a resident population of cells to divide and differentiate to form new muscle tissue69, and as we have already seen, BMPs can induce the formation of bone ectopically in muscle tissue29. These observa- tions have found resonance in stem-cell research with the result that many growth factors are important components in the differentia- tion regimes for both adult and embryonic stem cells70–72. In tissue engineering, the application of growth factors within biomaterials also represents a powerful tool for controlling cell differentiation and function. For instance, when murine muscle lacerations were treated by transplantation of muscle precursor cells within RGD- coupled alginate gels, recovery was greatly improved by the addition of hepatocyte growth factor (HGF) and fibroblast growth factor-2 (FGF-2)73. Already, growth factors feature in a handful of commer- cially available TE products (Table 1), one of which — Medtronic’s INFUSE — represents the field’s biggest financial success yet4. INFUSE is supplied with powdered recombinant human BMP-2, which is reconstituted in water and added to a collagen sponge immediately before use. Controlled-release strategies are frequently adopted to overcome the short half-life and residence of free growth factors in solution. For example, microspheres fabricated by double emulsion can release protein payloads from aqueous pockets within the particles, and can now be made to nanoscale dimensions using a single sur- factant74. Furthermore, in developmental pathways, different factors become active at different times, and growth factor release profiles that recapitulate these dynamics are likely to provide more lever- age over cell behaviour than those that apply growth factors indis- criminately. Materials schemes based on different degradation rates or diffusive properties of polymers have been designed with this in mind75,76 (Fig. 2). Although most efforts so far have concentrated on evaluating the effects of freely diffusible forms of growth factors in solution, most in fact function at interfaces in vivo, bound to ECM com- ponents or as part of membrane complexes77. Although concern undoubtedly arises over the cellular accessibility and activity of surface-immobilized proteins, even relatively simple tethering of a growth factor to a biomaterial matrix can elicit desired biological responses78,79. For example TGF-β1 covalently tethered to PEG not only retained its ability to stimulate matrix production in vascu- lar smooth muscle cells, but also did so significantly more than a comparable concentration of the soluble form of the protein80. Fixing growth factors covalently in place carries the added benefit of preventing internalization of growth-factor–receptor complexes by cells. More precise, site-specific couplings can be engineered through the use of recombinant proteins into which additional amino acids are introduced at the termini, for example Cys-tags or enzyme substrate sequences that lead to proteolytic release81,82. Systems for controlling the kinetics of growth factor release and presentation have shown potential for aiding blood vessel growth into scaffolds (Box 2, Fig. 3). More natural mechanisms of growth factor binding and release are also being pursued. In vivo, glycosaminoglycans (GAGs), mostly as components of proteoglycans, have key parts in growth factor activity, including sequestering them within the matrix, prevent- ing their degradation and presenting them to cell-surface recep- tors. GAGs are complex molecules with tissue-specific distribution and multiple physiological functions, but they share characteristic linear structures of repeating hexosamine-uronic acid disaccha- ride units83. Heparin, and heparan-, chondroitin-, keratin- and dermatan-sulphate GAGs (HS, CS, KS, DS, respectively) also have tightly regulated regiospecific sulphation patterns, which determine their specific interactions with proteins84,85. These interactions can be essential for the physiological effects of growth factors. FGF-1, for example, requires HS binding for dimerization and receptor activation86. Heparin has been widely incorporated into TE scaffolds to offer a slow release mechanism87,88 (Fig. 2), and CS in commer- cial products (Table 1) may perform a similar function, including modulating the activity of cell-derived signalling factors. simplifying biomolecules for use in biomaterials Few approved products include recombinant growth factors, but the enormous success of INFUSE shows the potential commercial viability of these material/growth factor combinations (Table 1): it attracted nearly US$700 million of sales in 2007 (ref. 4), an order of magnitude more than any of its competitor products. Furthermore, the sophisticated use of growth factors is likely to be important in advanced TE applications. For example, the patterning of growth factors within prefabricated scaffolds could aid the generation of heterogeneous tissues89. As already discussed for integrin-binding and protease-digestible proteins, growth-factor-mimicking thera- peutics where some of the growth factor function is condensed into relatively short peptide fragments, typically of 30–40 amino acids90,91, hold promise. Some of these peptides bind their respec- tive receptors with comparable affinities to recombinant growth fac- tors, and can trigger signal transduction and lead to appropriate cell responses. Although the concentrations required to elicit biological effects are variable, and in some cases exceed those of the native proteins by orders of magnitude, angiogenesis has been induced by one FGF-2 mimetic peptide at similar concentrations to recom- binant FGF-2 (ref. 91). This molecule contains two 15-amino-acid receptor-binding domains and a 9-amino-acid heparin-binding Rel eas e of > 1 g rowth fa ctor Mod es o f sp ati al pr es en ta tio n Scaold surface Dierent rates of diusion Dierent rates of polymer breakdown Loaded polymer and microspheres Loaded polymer coatings Release strategies Protease Cleavable peptide Cell- demanded release GAG sequestered Enhanced binding Free in solution Tethered: random orientation Tethered: specific orientation Use of spacer such as PEG Figure 2 | Presentation and release of growth factors from Te scaffolds. Anticlockwise, from top: growth factors within TE scaffolds may be loaded into polymers whose rate of degradation or diffusive properties can be modulated to tailor release rate, and which may be combined into systems releasing multiple factors with distinct kinetics75,76. The exposure of cells to different growth factors with time may therefore imitate developmental pathways and healing responses. An alternative to presenting growth factors in soluble form is to bind them to a surface in either random or specific orientations, with the possible use of a spacer molecule78–80. Non‑covalent associations with matrix components, particularly glycosaminoglycans (GAGs), can effect slow release and in some cases may potentiate binding to membrane receptors87,88. Cell‑demanded release is based on the presence of protease‑sensitive peptide sequences within the growth factor protein81,82. nmat_2441_JUN09.indd 463 12/5/09 15:34:18 © 2009 Macmillan Publishers Limited. All rights reserved 466 nature materials | VOL 8 | JUNE 2009 | www.nature.com/naturematerials review article NaTure maTerIalS doi: 10.1038/nmat2441 High LowMedium Structural synthetic mimics Anionic and phosphate groups Multi-domain peptide Sulphated and carboxylated dextran Enzyme-sensitive peptide crosslinkers Cryptic site peptides Sulphonated synthetic polymers Poly(glutamic acid) peptide Domain III Integrin binding sequence Fibronectin Structure and function of ECM molecules Protease Glycosylated synthetic polymers Proteoglycan aggregate Chondroitin sulphate, a GAGSulphated groups Proteoglycan Enzymatic cleavage sequences Amino acids Bone sialoprotein interaction R GD Cell behaviour Functional synthetic mimics Cell–cell adhesion Global response: including viability, adhesion and dierentiation Biotin–avidin crosslinking Polymer microarray Bone sialoprotein- derived peptide P C O O O O O O Synthetic polymer matrix Dimerized anity peptides Receptor binding Heparin binding Biotin Avidin Cell surface Protein complex Cell membrane E-cadherin Cytoskeletal actin Cell junctions Fibroblast growth factor 1 (FGF-1) bound as dimer FGF receptor 2 Bound heparan sulphate Collagen triple helices Hydroxyapatite crystals S O O O - SO O O- Oligosaccharides such as heparin oligomer - Poly(glutamic acid) sequences Integrin recognition sequences RGD IKVAV PHSRNYIGSR PDSGR EEEEEEEE E E EEEEEEEE ...GPQGIWQG... ...GP QG IWQG... e f Active fragment released Integrin binding, protease sensitivity Mineralization mediators Growth factors Protein binding a b d Carbohydrates c Proteins Figure 4 | Synthetic mimics of biological structures. Many characteristics of ECM macromolecules have been reproduced in simpler compounds with biologically inspired structures. a, Certain protein functions, including integrin binding for cell attachment and protease degradability, can be isolated to short amino‑acid sequences. These sequences can be combined with synthetic polymers or incorporated into complex peptides to enable cells to attach to or break down the material, respectively55,56,58–60,63,64. b, Some glycoproteins involved in bone mineralization, such as bone sialoprotein, possess runs of negatively charged amino acids. Peptides that incorporate these sequences, and synthetic polymers with negatively charged chemical groups, can display improved mineral‑nucleating activity120,121,128,129. c, Growth factor action has been demonstrated in peptides possessing receptor‑binding domains; heparin‑ binding sequences may also be included to aid growth factor sequestration90,91. Random peptide libraries have allowed the identification of peptides with affinity to particular receptors, and dimerization of these molecules in some cases can improve receptor binding and physiological response94. Growth factor action is sometimes potentiated through the actions of glycosaminoglycans (GAGs) such as the binding of heparan sulphate to the FGF‑1/FGF receptor 2 complex. This specific interaction can be achieved using short heparin oligosaccharides86. d, Furthermore, the protein‑binding function of ECM GAGs such as chondroitin sulphate can be mimicked by grouping sulphated oligosaccharides by polymerization, by sulphating natural carbohydrates such as dextran, or by sulphonating synthetic polymers84,99–102,130. Additionally, certain biological functions can even be supported by chemistries with no relation to biological structures. e, Whereas cell–cell adhesion occurs predominantly through the complex binding of cell surface cadherin proteins, biotin–avidin interactions have been used to artificially aggregate cells28. f, A range of responses, such as cell adhesion, viability and differentiation, can be differentially affected by particular synthetic substrates. These cell–material interactions can be assayed using high‑throughput screening of cells on polymer microarrays117–119. (Depictions of protein and peptides do not represent structures accurately.) nmat_2441_JUN09.indd 466 12/5/09 15:34:20 © 2009 Macmillan Publishers Limited. All rights reserved nature materials | VOL 8 | JUNE 2009 | www.nature.com/naturematerials 467 review articleNaTure maTerIalS doi: 10.1038/nmat2441 a non-adhesive polymer such as poly(hydroxyethyl methacrylate), which prevents cell migration between spots. Following cell culture, standard immuno histochemical techniques and microarray scan- ning can be performed. This provides a way of identifying poly- mers that support desired responses from specified cell types, for example those that promote differentiation of human embryonic stem cells117–119. Another approach has been to select chemical functionalities based on their resemblance to characteristic chemical features of particular ECMs. Earlier we described the bio-inspired use of sulphonate groups within hydrogels, mimicking their presence in GAG chains, which increased protein uptake102. This approach is well established in bone tissue engineering, where there exist many examples of materials incorporating anionic chemical moieties that improve mineral deposition, for instance by NaOH treatment of scaffold surfaces or by the incorporation of functionalized mono- mers such as methacrylated amino acids (GlyMA, SerMA, AspMA, GluMA)120,121. This practice stems from the observation that many glycoproteins involved in bone mineralization display a high pro- portion of negatively charged amino acids: for example, bone sialo protein (BSP-II) possesses two polyglutamic acid sequences, and osteopontin (BSP-I) contains a run of 10–12 aspartic acid resi- dues122. Phosphate groups also nucleate mineral, and although often delivered in soluble form in vitro (as β-glycerophosphate) can be incorporated covalently into scaffolds. Moreover, these chemical groups may also be instructive to cells. In a recent study, small defined chemical groups were incor- porated into PEG gels, and encapsulated human mesenchymal stem cells differentiated towards cells of those tissues that the functional groups chemically resembled54. Thus, those cells cultured in gels with charged phosphate groups increased expression of RUNX2 (CBFA1; an early bone transcription factor), produced a collagen- rich pericellular matrix, and synthesized osteopontin. Hydrophobic t‑butyl groups pushed cells towards an adipocytic (fat cell) lineage, demon strated by upregulation of the transcription factor PPARγ and the deposition of intracellular lipid deposits. It is unknown (and for practical purposes, arguably irrelevant) whether the role of the chemical modifications was to act directly on the cells, or to cause the preferential accumulation of particular cell-derived molecules, these molecules in turn providing behavioural signals to cells. An example of the latter mechanism in action is the ability of mineral deposits to sequester osteopontin, which improves cell adhesion and viability within phosphate-containing PEG gels123. Whatever the modes of action, the complexity of biomaterials could be massively reduced if the essential chemical character of ECM influences could be distilled into simple chemical functionalities. A summary of the various ways in which relatively simple molecules can mimic the molecular infor- mation within the ECM is given schematically in Fig. 4. Concluding remarks and perspectives The examples discussed herein demonstrate the importance of the extracellular environment in determining cell behaviour, and highlight the need for regenerative materials to provide cells with biological cues. Much is still unknown about the mechanisms by which tissues form and heal, yet already insights from developmen- tal biology and other biological disciplines are actively guiding the development of intelligent materials that work with nature’s own mechanisms of repair. These expanding possibilities raise the ques- tion of how much extrinsic physiochemical information is required to mobilize endogenous or transplanted cells into producing a complex tissue, and specifically, what minimum level of materials complexity is required for a given task. Evidently, a careful appraisal of the job in hand will reveal that the broader cost and treatment implications for any biomaterials approach vary with several interrelated factors including the form of the device, the mode of delivery, the nature of the cellular component and any regulatory implications. To elaborate briefly, injectable matrices help to tackle problems of surgical invasiveness whereas tissue engineering prod- ucts in sheet form (Table 1) confront problems related to nutrient supply by limiting diffusional distances. Moreover, materials that can recruit endogenous cells into scaffolds avoid the expense and difficulties associated with culture, storage and distribution of cells, not to mention immune considerations. Encouragingly, however, it is clear that comparatively simple materials in combination with an appropriate cellular component can support a high level of tis- sue organization14,60. The optimization of mechanical and struc- tural features of scaffolds and their potential to direct aspects of cell behaviour illustrates that functional sophistication is not necessarily synonymous with high manufacturing costs. A large number of commercially viable products for connective tissues are based on purified ECM components such as collagens and hyaluronic acid (Table 1), representing a relatively generic ECM backdrop conducive to the activities of differentiated cells. Imposing a tissue-specific identity on stem cells in many cases is likely to require more specific influences, within materials if not during cell culture. These more advanced biomaterial approaches are just beginning to trickle through product-development pathways, but the runaway suc- cess of INFUSE4 demonstrates the potential impact of schemes that make use of growth factor activity. It is promising that the outcome of growth factor administration can be improved enormously with the use of technically simple slow-release schemes, such as delivery using polymers. Such considerations may prove critical for the resolu- tion of complex tissue engineering challenges such as that of vascu- larization (Box 2). However, the generation of thick or heterogeneous constructs, and even complex organs, will require further innovations in biomaterials research. Interest is also growing in the exciting pos- sibility of using simple chemistries to influence cell behaviour, and in the development of a range of therapeutics with intrinsic or modulat- ing growth factor activity, including designer carbohydrates. Several laboratories in their own ways are actively pursuing simple but effec- tive solutions to tissue engineering problems, such that the ideal of structurally simple, yet functionally complex, biomaterials is becom- ing a plausible possibility for the near future. 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