Controlled drug delivery in tissue engineering

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ews 60 (2008) 229–242www.elsevier.com/locate/addr

Advanced Drug Delivery Revi

Controlled drug delivery in tissue engineering☆

Marco Biondi, Francesca Ungaro, Fabiana Quaglia, Paolo Antonio Netti ⁎

Interdisciplinary Research Centre on Biomaterials (CRIB), University of Naples Federico II, Naples, Italy

Received 1 August 2007; accepted 9 August 2007Available online 11 October 2007

Abstract

The concept of tissue and cell guidance is rapidly evolving as more information regarding the effect of the microenvironment on cellularfunction and tissue morphogenesis become available. These disclosures have lead to a tremendous advancement in the design of a new generationof multifunctional biomaterials able to mimic the molecular regulatory characteristics and the three-dimensional architecture of the nativeextracellular matrix. Micro- and nano-structured scaffolds able to sequester and deliver in a highly specific manner biomolecular moieties havealready been proved to be effective in bone repairing, in guiding functional angiogenesis and in controlling stem cell differentiation. Althoughthese platforms represent a first attempt to mimic the complex temporal and spatial microenvironment presented in vivo, an increased symbiosis ofmaterial engineering, drug delivery technology and cell and molecular biology may ultimately lead to biomaterials that encode the necessarysignals to guide and control developmental process in tissue- and organ-specific differentiation and morphogenesis.© 2007 Elsevier B.V. All rights reserved.

Keywords: Tissue engineering; Drug delivery; Biomaterials; Growth factors; Scaffold

Contents

1. Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2302. Extracellular matrix mimicry as guideline for scaffolds design . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2303. Tissue engineering scaffolds as controlled release matrices . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 232

3.1. Interspersed signals . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2323.2. Immobilized signals . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2343.3. Signal delivery from cells . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 234

4. Delivery systems for proteins of potential interest in tissue engineering . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2364.1. Continuous delivery . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 236

4.1.1. Non biodegradable systems . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2364.1.2. Biodegradable systems . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 236

4.2. On–off delivery . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2374.2.1. Programmed delivery systems . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2374.2.2. Triggered delivery systems . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 237

5. The issue of delivery system integration in three-dimensional scaffolds . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2376. Conclusions . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 238References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 238

Abbreviations: bFGF, basic fibroblast growth factor; BMP, bone morphogenetic protein; BSA, bovine serum albumin; CASD, computer-aided scaffold design;DS, delivery systems; ECM, extracellular matrix; EGF, epidermal growth factor; EVAc, ethylene-vinyl acetate copolymers; GF, growth factor; HBDS, heparin-baseddelivery systems; NT-3, neurotrophin-3; PA, peptide amphiphile; PCL, poly(ɛ-caprolactone); PDGF, platelet derived growth factor; PEG, poly(ethylene glycol); PEO, poly(ethylene oxide); PLA, poly(lactide); PLGA, poly(lactide-co-glycolide); POE, poly(ortho esters); PTH, parathyroid hormone; SFF, solid free-form fabrication; TE, tissueengineering; TGF-β1, transforming growth factor-beta1; VEGF, vascular endothelial growth factor.☆ This review is part of the Advanced Drug Delivery Reviews theme issue on “Emerging Trends in Cell-Based Therapeutics”.* Corresponding author. Tel.: +39 817682408; fax: +39 817682404.

E-mail address: [email protected] (P.A. Netti).

0169-409X/$ - see front matter © 2007 Elsevier B.V. All rights reserved.doi:10.1016/j.addr.2007.08.038

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230 M. Biondi et al. / Advanced Drug Delivery Reviews 60 (2008) 229–242

1. Introduction

Tissue engineering (TE) aims at the repairing and restor-ing damaged tissue function employing three fundamental“tools”, namely cells, scaffolds and growth factors (GFs)which, however, are not always simultaneously used [1,2].On the other hand, summoning recent experimental and cli-nical evidences indicate that the success of any TE approachmainly relies on the delicate and dynamic interplay amongthese three components and that functional tissue integrationand regeneration depend upon their sapient integration [3,4].Future generation of scaffolds will have to provide not onlythe adequate mechanical and structural support but also haveto actively guide and control cell attachment, migration, pro-liferation and differentiation. This may be achieved if thefunctions of scaffold are extended to supply biological signalsable to guide and direct cell function through a combinationof matricellular cue exposition and GF sequestration anddelivery [2,5]. Therefore an ideal scaffold should possess athree-dimensional and well defined microstructure with aninterconnected pore network, mechanical properties similar tothose of natural tissues, be biocompatible and bio-resorbableat a controllable degradation and resorption rate as well asprovide the control over the sequestration and delivery ofspecific bioactive factors to enhance and guide the regener-ation process [6,7].

Recent advances in micro- and nano-fabrication technologiesoffer the possibility to engineer scaffolds with a well definedstereoregulated architecture providing a control of cellular spa-tial organization, mimicking the microarctitectural organizationof cells in native tissues [6,8–13]. Furthermore, by combiningmaterial chemistry and processing technology, scaffold degra-dation rate can be tuned to match the rate of tissue growth in sucha way that the regenerated tissue may progressively replace thescaffold [14–16]. Enhancing further the functionality of thesealready complex matrices by encoding in them the capabilityto expose an array of biological signals with an adequate doseand for a desired time frame, represents the major scientific andtechnological challenge in tissue engineering today. Bolus ad-ministration of GFs would not be effective in these cases sincethey rapidly diffuse from the target site and are readily enzy-matically digested or deactivated. Moreover, local delivery andprolonged exposition of the bioactive molecules is necessary tominimize the release of the agent to non-target sites, and supporttissue regeneration which normally occurs in long time frames[17]. Thus, it has been soon realized that by integrating con-trolled release strategies within scaffoldingmaterials may lead tonovel multifunctional platforms able to control and guide thetissue regeneration process [18–22]. Through the recapitulationof the spatial and temporal microenvironments presented bynatural extracellular matrix (ECM), it is hoped to successfullyguide the evolution of the construct towards neotissue formation,inducing on-demand different pathways to cell response. In thisperspective, TE can be viewed as a special case of controlleddrug delivery in which the presentation of bioactive molecules isfinely tuned to dynamically match the needs of the ingrowingtissue.

The control over the regenerative potential of TE scaffoldshas dramatically improved in recent years, mainly by using drugreleasing scaffolds or by incorporating drug delivery devicesinto TE scaffolds [17,19,23]. For example, on-demand respon-sive matrices based on enzymatically-triggered release of GFshave been realized by introducing enzyme-cleavable linkers forcovalent interaction between the released molecule and a bio-active protein [8]. Furthermore, potent morphogenetic factorshave been loaded in polymeric depots and included into variousbiomaterials to enable a sustained and controlled point sourcerelease while preserving bioactivity as reviewed extensively inthe literature [19–22]. Despite the impressive enhancement intissue guidance and regeneration offered by GF releasing scaf-folds, several challenges have yet to be broadly resolved. Theseinclude the tight control over time and space of tiny quantitiesof multiple biomacromolecular factors and of their gradientswithin the interstitial space of the scaffold as well as at thescaffold-tissue interface. Moreover, there is a paucity of studiesregarding the effective dose in the local microenvironment, themagnitude of the spatial and temporal gradients and the de-velopment of technological strategies to integrate and positiondrug delivery devices with a submicrometric spatial resolutionwithin the scaffolds.

In this review we will first summarize the complex processesof cell guidance occurring within native ECM along with themost updated strategies to design biomimetic scaffolds able torecapitulate in part these processes. A synthetic overview of themost promising approaches in controlling the release of therelevant factors in TE will follow. Finally, the main challengesto design novel scaffolds with time and space orchestratedexposure of biomacromolecular moieties will be presented andcritically discussed.

2. Extracellular matrix mimicry as guideline for scaffoldsdesign

ECM, the natural medium in which cells proliferate, differ-entiate and migrate, is the gold standard for tissue regeneration[24]. Cell-ECM interaction is specific and biunivocal. Cellssynthesize, assembly and degrade ECM components respond-ing to specific signals and, on the other hand, ECM controlsand guides specific cell functions. The continuous cross-talkbetween cells and ECM is essential for tissue and organdevelopment and repair, providing both a structural guidance(i.e. directional cell migration) and cell guidance at a molecularlevel (i.e. signaling molecule delivery).

ECM is a highly organized dynamic biomolecular environ-ment in which many proliferation–adhesion–differentiationmotifs, governing cell behaviours, are continuously generated,sequestered and released, inducing matrix synthesis and degra-dation (Table 1). These motifs are locally released according tocellular stimuli, generally occurring upon degradation of theadhesion sites binding them to the ECM [25]. Cells areattached to ECM through molecules belonging to the integrinfamily [26] and recognize specific amminoacid sequencesthrough cell surface receptors. Integrin receptors are recruitedin microdomains of cell membrane, and in these areas integrins

Table 1Main molecular components of ECM and their role

Component Description Function Location Ref.

Collagens Fibrillar protein forming ECM backbone Resistance to tensile strength Ubiquitous [148,149]Matrix scaffoldingCell–cell interactionCell–ECM interactionModulation of ECM morphologyModeling the framework of connective tissues

Proteoglycans Carbohydrate polymers composed ofa polypeptide backbone covalently attached toglycosaminoglycan chains

Filler substance between cells Ubiquitous [150]Binding to cations/waterTransport of small molecules in ECMResistance to conprhessive stressesCell adhesion, migration and proliferationLigand/receptors of signaling molecules

Hyaluronan Negatively charged glycosaminoglycan polymer Cell proliferation Ubiquitous [151,152]Cell differentiationCell migrationTransport of small molecules in ECM

Laminins Glycoproteins of basal lamina Development and maintenance of basement membranes Basementmembranes

[153]Receptor-mediated cell attachmentCell signaling/migration

Fibronectin High molecular weight glycoprotein binding toECM components

Cell adhesion Cell surfaces [154]Cell migrationCell proliferationMatrix adhesionFibroblast activation

Elastin Hydrophobic, cross-linked insoluble protein Tissue resilience and elasticity Blood vessels [155]LungsLigamentsSkinBladder

Growth factors Proteins associated to ECM or heparin sulphate able toinduce cell migration, proliferation, differentiation,in soluble form, upon activation of latent forms

Information processing Ubiquitous [33,156]Cell signalingECM synthesis/remodeling

231M. Biondi et al. / Advanced Drug Delivery Reviews 60 (2008) 229–242

communicate with structural and signaling molecules influ-encing transport, degradation and secretion of ECM mole-cules, endocytosis and cellular fate [27,28]. Moreover, solid-state, structural ECM molecules act as reservoirs for secretedsignaling molecules for their on-demand release [29,30].Besides insoluble factors and proteins presented on the surfaceof adjacent cells, the molecular cues that mainly define themicroenvironment consist of soluble macromolecules, such asGFs [31].

GFs are protein molecules specific for intercellular andcell-ECM signaling, which are involved in ECM dynamicproperties through specific surface receptors, driving GFsregulatory activity [31,32]. GFs are locally secreted by ECM,in which they are stored in insoluble/latent forms throughspecific binding with glycosaminoglycans (e.g. heparins), andcan elicit their biological activity once released. During tissuemorphogenesis the presence of soluble GFs guides cellularbehaviours, thus governing neo-tissue formation and organi-zation. The sequestration of GFs within ECM in inert form isnecessary for rapid signal transduction, allowing extracellularsignal processing to take place in time frames similar to thoseinside cells. Moreover, concentration gradients of GFs play amajor role in ECM maintenance and equilibrium becausethe gradients direct cell adhesion, migration and differentia-tion deriving from given progenitor cells and organize pat-

terns of cells into complex structures such as vascularnetworks and nervous system [33–35]. Thus, spatial patternsin tissues are dictated by both the architectural features ofECM and concentration profiles/gradients of diffusible bioac-tive factors [36].

Recent research in biomaterial science has been driven bybiomimicry-inspired design of materials to recreate the naturalthree-dimensional architecture. Several micro- and nano-fabrication strategies have been applied in an attempt to mimicthe spatial distribution of the fibrillar structure of ECM, whichprovides essential guidance for cell organization, survivaland function [9–13]. These technologies include gas foaming,solid free-form fabrication (SFF) (3D printing, 3D plotting),molecular and nanoparticulate self assembly, electrospinning,molecular and nano-templating [6,9,12,13,37]. Albeit theinfluence of scaffold microarchitecture and stereomorphologyon cell function and guidance has been proved in severalsystems and with different cell types, the underlying mechan-isms by which cells recognize and decode topological infor-mation are still unclear. A wide variety of biodegradable andbiocompatible polymers have been processed to fabricatestereoregulated scaffolds, including synthetic polymers, suchas poly(lactide) (PLA), poly(glycolide) and their copolymerspoly(lactide-co-glycolide) (PLGA), poly(ɛ-caprolactone)(PCL), and natural polymers, such as collagen, protein, and


fibrinogen [14–16] (Table 2). Most of the technologies used todate suffer from the limitation that the scaffolds are preformedand cell have to be loaded within the interstitial spaces whichare often smaller than the cell size.

3. Tissue engineering scaffolds as controlled releasematrices

Future generations of TE scaffolds with extended function-ality and bioactivity require an increased integration with celland molecular biology, to identify novel design parameters andbio-inspired design approaches (Fig. 1). Synthetic bio-inspiredECM should broadcast specific cellular events, such as therecruitment and the enhancement of migration of peripheral hostcells into the scaffold, or should guide morphogenetic processestaking place within its interstices through the fine tuning ofspatial and temporal gradients of growth and morphogeneticfactors [38–42].

The local concentration and the spatio-temporal gradients ofa molecule depend upon a delicate balance between thetransport properties of the scaffold, the binding and degradationrate of the molecule and its generation rate. Once transportmechanisms and biological decay time constant of the bioactivemolecule are known, it is virtually possible to engineer anycomplex static or dynamic gradient distribution within thescaffold by including artificial reservoirs able to deliver the

Table 2Materials for tissue engineering scaffolds

Material

Naturally-derived Collagen-based scaffolds

Hyaluronic acid (HA) and derivatives

Collagen-HA gelsChitosan

Fibrin

Gelatin

AlginateSynthetic Poly(glycolic acid) (PGA)

Poly(lactic acid) (PLA)Polylactide-co-glycolide (PLGA)

Poly(ɛ-caprolactone) (PCL)

Polyethylene glycol (PEG)

Oligo(poly(ethylene glycol) fumarate) (OPF)

Inorganic Tricalciumphosphate (TCP)Hydroxyapatite (HA)

Semi-synthetic Cross-linked thiolated HA

Esterified hyaluronan (HYAFF® derivatives)

relevant biomacromolecules at a predefined rate. However, themagnitude of the gradients and the optimal time frame to elicitthe desired cell response are yet unknown. Significantadvancement in scaffold design can be achieved through adeeper understanding of the quantitative aspects of the influenceof amount of morphogens and their gradient on cell fate.

One of the possible attempts to control molecular microen-vironment is to use a TE scaffold as a controlled releaseplatform. This can be achieved by the incorporation of signalingmolecules in three-dimensional scaffolds through their simpledispersion in the matrix, or their immobilization by electrostaticinteractions with and covalent bonding to the scaffold. A gene-mediated approach is also feasible by introducing into targetcells nucleic acid encoding for a specific protein signal inducingtissue regeneration. In this way, each cell can act as a singlesource point for release of the signaling proteins.

3.1. Interspersed signals

Signaling molecules can be integrated within scaffolds bysimply interspersing them in the matrix. Although this methodpresents several shortcomings, it has been widely used in theliterature. Most of these approaches are carried out by hydrogel-based scaffolds in such a way that the hydrogel actssimultaneously as a scaffold and a controlled delivery platform.Both naturally-derived (collagen, fibrin, chitosan) [14,15] or

Relevant features and application Ref.

Soft tissue repair [16,157,158–160]Cell differentiationCapillary engineeringDermis engineeringVascularized adipose tissueRegeneration of skin, cartilage [161]Patterning of cell growthControl of vascular sprouting [162]Chitosan microsphere-integrated scaffold [163,164]Cartilage engineeringVessel engineering [165,166]Release of fibroblastsTrachea engineering [167,168]Bone engineeringVascular engineering [100]Musculoskeletal tissue [8,169]engineeringCartilage regenerationFibrovascular engineeringSkin engineering [170]

Bone formation [171]

Cartilage engineering [172]GF release from gelatin microspheresBone substitute [173,174]

Neurite growth and support [175,176]Vocal fold repairCartilage engineering [160]

Fig. 1. The concept of multifunctional TE scaffolds based on controlled delivery technology. A hydrogel or solid state three-dimensional scaffold (A) releases GFs,either encapsulated in controlled delivery systems, or dispersed/tethered within the scaffold, according to predetermined spatial gradients and with controlled kinetics,occasionally with retarded delivery onset (B). The broadcasting of the appropriate molecular signals towards the surrounding defective tissue induces the desiredcellular responses (C), which are triggered by the spatio-temporally controlled presentation of the proper tissue-inductive signals.


synthetic hydrogels (poly(ethylene glycol) (PEG)) have beenused [14,16]. Natural materials possess innate capacity ofcellular interaction and undergo a mainly cell-related degrada-tion profile, while synthetic materials lack the cellular/biomolecular recognition but can be more readily manipulatedin terms of macro- and microscopic properties [8].

Under particular conditions, cells and other bioactive entitiescan be safely encapsulated in hydrogels before gelation, thestructural parameter that mainly controls transport propertiesbeing the cross-linking density and stability [43,44]. Thedispersed factors move through the mesh network of thecrosslinked gel by a combination of diffusion and degradationmechanisms. Particularly, when the hydrodynamic diameter ofthe diffusing species is smaller than the average hydrogel meshsize, diffusion mechanisms prevail. This leads to a fast, notsustained release (hours to days) of the dispersed molecules,which is not very useful for a TE approach. On the other side,when the molecular size approaches hydrogel mesh size, therelease is mainly controlled by the degradation of the polymerbackbone, which can occur either by hydrolytic or enzymatic(i.e. on cell demand) mechanisms [45].

A variety of synthetic and natural polymers have beenutilized for the design of hydrolytically degradable hydrogelsin which chemical or physical cross-linking offers the possi-bility to control the diffusion of solubilized hydrophilicmacromolecules [46–48]. An important formulation challengewhen fabricating hydrogels for protein delivery is the choice ofthe cross-linking method, which must not involve steps withpotentially adverse effects on protein stability. Naturally-derivedhydrogels such as collagens, hyaluronic acid and derivatives arefrequently used due to their well-known biocompatibility. Forexample, collagen-based scaffolds can induce transgene expres-sion and physiological improvements for bone regeneration [49]

and wound healing [50]. Furthermore, collagen modified withpoly(L-lysine) has been shown to promote aspecific interactionsbetween plasmid molecules and collagen, resulting in plasmidbinding and release [51].

Among synthetic hydrogels, PEG-based materials are widelyapplied in TE. Release mechanism and degradation rate can betailored through chemical modifications by inserting unitsaffecting PEG functionality, such as fumarate, lactic acid,caprolactone, hyaluronic acid units [52–54]. Pseudo-zero-orderrelease kinetics can be attained by modulating the degradationrate of cross-links and cross-link density by inserting additionaldouble bonds [55], or through the optimization of cross-linkingagent amount [56,57]. Proteolytically-sensitive PEG-based net-works have been engineered by the group of Hubbell byencoding signals able to finely control the release of bioactiveagents based on a cell-demanded logic [58]. These syntheticanalogs of ECM can thus represent a first step towards thereproduction of the dual-reciprocity mechanism (i.e. cell-materialcross-talk) occurring in a native ECM.

Synthetic solid biodegradable materials have been tested fordrug delivery in TE, especially for hard-tissue repair. However,also the fabrication of protein-loaded solid scaffolds posesserious issues regarding protein leaching and stability [19],such as: i) the use of organic solvents during the manufac-turing processes; ii) protein elution occurring during scaffoldprocessing (e.g. hydrosoluble porogen leaching processes);iii) exposure of the protein to high temperatures (e.g. meltprocessing); iv) generation of organic-aqueous interfaces duringscaffold processing (e.g. scaffolds made of lipophilic syntheticbiodegradable polymers). Thus, direct encapsulation of proteinsin solid scaffolds should be preferentially carried out under mildtechniques, such as gas foaming and electrospinning, eventuallycombined with particulate leaching. Despite these limitations,


scaffolds made of PLGA have been successfully used toengineer tissues. Release through PLGA occurs by a combineddiffusion-degradation mechanism which leads to the progres-sive generation of acids, particularly in the bulk regions of thescaffold [59], and have been produced for the prolonged releaseof vascular endothelial growth factor (VEGF) for bone TE [60]and vascular bed generation [61], or a plasmid encoding forplatelet-derived growth factor (PDGF) in vascular induction[62].

The few examples described above show that scaffolds inwhich the bioactive agent is simply dispersed may not offer thenecessary control over release kinetics and extent. Therefore,new designing approaches have been exploited to provide acontrol over spatial and temporal release pattern.

3.2. Immobilized signals

Polymer scaffolds can be modified to interact with signalingmolecules, thereby hindering their diffusion out of the polymerplatform, thus prolonging their release. Signal immobilizationcan occur through reversible association with the scaffold (i.e.binding/de-binding kinetics), irreversible binding to the poly-mer. Alternatively, signals can be released upon degradation of alinking tether or the matrix itself, which immobilize themolecule within the scaffold. The number of binding sites, theaffinity of the signal for these sites, and the degradation rate ofthe scaffold are key determinants of the amount of bound signal,as well as the release profile [40,63].

In the case of GFs, the most common approach to improverelease kinetics of the immobilized molecule relies on the use ofheparin-immobilized scaffolds. Actually, heparin grafted on thesurface or chemically bound to the polymer can interact withheparin-binding GFs [64–70]. Heparin-based delivery systems(HBDS) have been largely employed to control GF concentrationwith fibrin scaffolds, in which a synthetic linker peptide, capableof binding heparin, is covalently attached to fibrin [64,66,67].Conjugating capacity and release rate were found to be dependentupon the number of binding sites, the affinity of factors towardsbinding sites and the degradation rate of the scaffold. HBDS forthe controlled release of neurotrophin-3 (NT-3) were fabricatedusing a linker peptide containing a Factor XIIIa substrate tosequester heparin within fibrin gels [67]. The authors showed thatheparin not only binds NT-3, hindering its diffusion, but alsoallows an active release mechanism, which is triggered by cell-associated enzymatic activity. In so doing, release of NT-3 wasextended for 9 days, and the neural fiber density was increased inspinal cord lesions. Also heparinized cross-linked collagens for invivo endothelial cell seeding have been studied with respect to GFbinding and release [65,70]. Collagen matrices were modifiedwith heparin for binding and release of basic fibroblast growthfactor (bFGF), by a conjugation reaction between carboxylgroups of heparin and amino groups of collagen [65]. In asystematic study, Nillesen et al. [70] prepared and characterizedfive porous scaffolds consisting of collagen, collagen withheparin, and collagen with heparin plus one or two GFs (bFGFand VEGF). The scaffolds obtained by collagen–heparinconjugation and GF incorporation displayed the highest density

of blood vessels and most mature blood vessels after subcutane-ous implantation in rats. Also synthetic scaffolds based onheparin–conjugated PLGA fabricated by a gas-foaming/salt-leaching method showed the ability to sustain bFGF release over20 days and promote blood vessel formation in vivo [68]. Morerecently, a heparin-conjugated PLGA scaffold for the sustaineddelivery of bone morphogenetic protein (BMP)-2 was used toenhance ectopic bone formation [69]. The amount of heparinconjugated to the PLGA scaffolds could be increased up to 3.2-fold by using scaffolds made from star-shaped PLGA, ascompared to scaffolds made from linear PLGA.

Immobilization of GFs within the scaffold can be alsoachieved by their covalent conjugation to the polymer [71–73].Covalent conjugation of GFs was intuitively thought not feasiblesince chemical bond could negatively affect their biologicalactivity. However, if appropriately designed, covalent conjuga-tion, or tethering, of GFs, have been proven a valuable strategy toretain GFs for longer time periods at the delivery site, offering animportant control over the amount and spatial distribution of thesemolecules in solid matrices. Whether covalently conjugated GFsact directly in the immobilized form, mimicking the heparin-bound complex, or act upon release occurring via hydrolyticcleavage of the tether is still to be ascertained. Early studiesdemonstrated that epidermal growth factor (EGF) covalentlytethered to aminosilane-modified glass via star poly(ethyleneoxide) (PEO) could elicit DNA synthesis and cell responses ofprimary rat hepatocytes, whereas the simple physical absorptionof EGF on modified glass was not effective [71]. Similarly,tethering transforming growth factor-beta1 (TGF-β1) to adhesiveligand-modified glass surfaces resulted in a significant increase inECM production over the same amount of soluble TGF-β1 [73].Using carbodiimide chemistry, BMP-2 was directly immobilizedon silk fibroin films [72]. Human bone marrow stromal cellscultured on unmodified silk fibroin films in the presence ofosteogenic stimulants exhibited little if any osteogenesis, whereasthe same cells cultured on BMP-2 decorated films in the presenceof osteogenic stimulants differentiated into an osteoblasticlineage. In another study, covalent attachment of VEGF andRGD-containing synthetic oligopeptides to PEG hydrogels couldgenerate complete vascularization of the construct by a cell-demanded release of the angiogenic factor [74].

These studies demonstrate that signaling by immobilizedGFs may be more potent than signaling by soluble GFs directlyinterdispersed within the scaffold. In particular, presentation ofGFs by covalent immobilization on scaffold surface may permita greater control of their temporal and spatial availability in theextracellular environment [75]. Nevertheless, the immobiliza-tion strategy must consider protein structure and active regiontopology when designing suitable delivery system (DS) in orderto maximize GFs bioactivity. Furthermore, some signals takesadvantage of sustained release, while others benefit from directattachment to the biomaterial substrate [76].

3.3. Signal delivery from cells

An alternative and more sophisticated approach to overcomeissues related to tissue regeneration (i.e. local and controlled

Fig. 2. Protein release from non-biodegradable (A), soluble/biodegradable (B) and pulsatile (C) delivery systems. Continuous, delayed and pulse-like delivery may beachieved with non-biodegradable and soluble/biodegradable delivery systems (D). On–off delivery (single or multi-pulses) may be achieved with pulsatile (pre-programmed or triggered) delivery systems (E).


delivery of GFs) and to elicit the desired biological responseswithin the scaffold relies on the use of nucleic acid-releasingsystems. In principle, nucleic acids containing a sequenceencoding for specific proteins can be introduced into targetcells, which are thus prompted to produce the desired proteins.Alternatively, oligonucleotides can be used to return abnormalgene expression to a certain state in antisense and interferencetherapies based on silencing RNA [77]. In this way, cellsgenetically induced to secrete proteins may act as point-sourceDS, allowing a prolonged and more specific effect.

For a successful gene-mediated TE approach, syntheticoligopeptides containing the adhesion site of fibronectin (theRGD sequence) have been used [36,78]. The nanoscaledistribution of the oligopeptides chemically coupled to adhesionsubstrates was found to mediate the efficiency of gene delivery[36]. Gene expression levels increase with increasing oligopep-tide density (i.e. stiffness of adhesion), which enhances theability of cells to internalize plasmid DNA. The use ofbiomaterial-based devices modified with specific cell-adhesionmolecules can maximize the population of stimulated cells [78].In perspective, coupling oligopeptides containing the RGDsequence to protein-loaded DS may improve the cellularresponse to GFs by exposing cells to genes according to apredetermined scheme, thus activating the tissue-inductive cellmachinery.

Synthetic gene-activated matrices, loaded with plasmid, mayalso play a prominent role as cell-activating scaffolds for TE.Time-controlled release of the plasmid encoding for tissue-inductive PDGF from porous PLGA matrices lead to matrixdeposition and vascular bed formation [62]. Similarly, thedispersion of plasmid DNA encoding for both angiogenic and

osteogenic factors within PLGA scaffolds resulted in a gene-activated cell recruitment from peripheral tissue promotingosteogenesis [79]. Bioengineered tissues secrete recombinantproteins and act as long-term DS when implanted in vivo.Myoblasts retrovirally transduced to locally secrete recombi-nant VEGF induced the regrowth of a functional capillary bedin the bulk of a bioengineered tissue substrate and in theadjacent muscle ischemic tissue [80]. More recently, geneticallyactivated rabbit bone marrow stromal cells engineered toexpress BMP-4 in a porous PCL scaffold effectively promotednew bone tissue formation [81].

The concept of gene-mediated protein expression has beenput forward by immobilizing DNA at the pericellular level, thustriggering molecular signal broadcasting from cells, probablydue to high local concentration of DNA in the cellular niche[82]. Combining GF delivery with covalent attachment of DNAable to dictate protein release from the cells may improve tissueresponse. Actually, dual delivery of BMP-2 and a plasmidencoding the same protein may induce a feedback mechanismby which the transcription efficacy of the plasmid was furtherincreased, thus opening the way to possible pathways eludingviral gene transfer [83].

The results of combined GF and gene delivery suggest theopportunity of a sequential release of multiple signalingmolecules, allowing the recapitulation of tissue formationsteps at predetermined time intervals and/or after inductiontimes. However, the synchronization between the intracellulartransport of gene material and the activation of proteinexpression by cells is still far from being optimized. Currentunderstanding suggests that the properties of the pericellularspace, the delivery method and the biomolecule structure must


be carefully controlled to achieve a synergistic effect able tolead to an overall improvement of the therapeutic approach.Probably, cellular signaling and genetic manipulation shouldwork in concert to promote a full mimicry of the naturalsequences governing tissue regeneration.

4. Delivery systems for proteins of potential interest intissue engineering

Drug delivery technologies can be of help in designingbioactivated scaffolds in which low or high molecular weightmolecules should be released in a specific area at pre-programmed rates [17,18,84]. First of all, a DS can offer to its“protein cargo” adequate protection from inactivation occurringin biological environments and guarantee the preservation ofbioactivity during the whole release duration [85]. On the otherhand a fine tuning of release rate can be realized by regulatingplatform composition, shape and architecture. DS offering atime-control of the delivered dose can be useful to trigger off therelease of a bioactive molecule and maintain a specificconcentration for extended duration. Furthermore, this strategygives the opportunity to deliver more than one protein atdifferent pre-programmed rates according to the needings of aspecific application.

DS can be designed in different shapes (particles, implants),architectures (reservoirs, matrices) and made with differentbiodegradable and non-biodegradablematerials offering a tunablecontrol over release rate. Twomain possible rates are feasible thatis continuous delivery and pulsatile delivery (Fig. 2). In thefollowing we describe a number of DS for proteins of potentialinterest in TE highlighting the type of control over release rateoffered.

4.1. Continuous delivery

4.1.1. Non biodegradable systemsPure diffusion-controlled systems based on non-biodegrad-

able polymers, such as ethylene-vinyl acetate copolymers(EVAc) and silicones, have been firstly tested/used for thecontrolled release of drugs. At present, the pharmaceutical useof EVAc for the controlled release of proteins is relatively rare,even if its potential has been long investigated and a particularattention has been devoted to GF release [86–88]. In suchsystems, protein transport out of the device is driven by aconcentration gradient and limited by the presence of aninsoluble polymeric matrix which regulates drug diffusion.Mass transport occurring through polymer chains or poresis the only rate-limiting step of the process. Reservoir ormatrix systems can be designed to respectively achieve zero-or first-order release kinetics, with different biologicalimplications. In this perspective, EVAc-based systems maybe of help in applications where the effect of GF releasekinetics within the scaffold on tissue regeneration processneeds to be highlighted.

Recent advances in the field of micro- and nanotechnologyhas given a new strength also to the application of silicon tubingin protein delivery. Actually, novel delivery and sensing silicon-

based platforms for long-term integration of cells may beachieved, forming the so-called ‘nanoporous micromachinedbiocapsules' [89]. These systems are specifically intended forencapsulation of pancreatic islet cells, able to release insulin, orother cells of interest. In perspective, applications other thanpeptide and protein delivery may involve the restoring of organfunctions.

4.1.2. Biodegradable systemsEven if EVAc and other non-biodegradable polymers are still

investigated as protein DS, current studies have been directedtowards the development of soluble/biodegradable systemsrequiring no follow-up surgical removal once the drug supply isdepleted. Amongst synthetic biodegradable polymers, thermo-plastic aliphatic polyesters like PLA and PLGA have generatedtremendous interest due to their excellent biocompatibility aswell as the possibility to tailor their in vivo life-time from weeksto years by varying composition (lactide/glycolide ratio),molecular weight and chemical structure (i.e. capped anduncapped end-groups) [90]. Protein encapsulation within PLGAmicro- and nano-carriers is regarded as a powerful tool toprotect the biological activity of generally labile macromolec-ular therapeutics and sustain their release over long time frames[91]. Different PLGA formulations for protein release arealready on the market (Lupron Depot®, Sandostatin LAR®Depot, Nutropine Depot® and Zoladex®) and several examplesof successful protein and GF delivery through PLGA micro-spheres are reported in the literature [92–102]. PLGA-basedparticles can be engineered in terms of composition, size (i.e.,microparticles or nanoparticles), size distribution and morphol-ogy to tailor the release rate on the specific application, and theirsurface functionalized to enhance their interaction with cell andtissues [103–105].

Drug-containing solution of PLGA copolymers in biocom-patible organic solvent have been proposed as in situ formingDS [106–107]. This technology (Atrigel®) is used in Eligard®(QLT Inc.), a leuprolide acetate-containing formulation for thetreatment of prostate cancer able to sustain protein release fromone to six months. Analogously, thermally induced gellingsystems based either on PLGA-PEG-PLGA or PEG-PLGA-PEG triblock copolymers have also been used as in situ formingprotein delivery platforms [106] and recently tested for thesustained release of interleukin-2, human growth hormone andinsulin [108–110]. Along with in situ forming biodegradablepolymers matrices, another class of biodegradable polymers isattracting research attention as protein DS due to their peculiarchemical–physical properties. The new generation of poly(ortho esters) (POE) have evolved through a larger number offamilies spanning from injectable viscous biomaterials, wherethe protein can be directly incorporated by simple mixing,without the use of heat or solvents, to a low melting temperaturepolymer (POE IV) that can be extruded at temperaturescompatible with protein biological activity [111,112]. Bothsemi-solid and solid extruded POE have been proved as highlyflexible and programmable matrices for controlled proteindelivery [113–115]. These systems are particularly relevant inTE since the polymer formulation can be easily adopted with the


most promising micro- and nanotechnologies adopted to fabri-cate scaffolds such as 3D printing and SFF.

A quite recent class of biodegradable polymers currentlyinvestigated for protein delivery are polyanhydrides, character-ized by an hydrophobic backbone carrying hydrolytically labileanhydride linkages [116]. Differently from PLGA copolymers,polyanhydrides are believed to undergo predominantly surfaceerosion providing a better and easier control over the proteinrelease kinetics through the material formulation [117].Polyanhydrides can be processed in several usable form, suchas particulate (i.e. microspheres, nanoparticles and beads) andmatrix systems (i.e. implant, films, surgical paste and sheets)[116,117] and have been proven to preserve the biologicalstability of protein therapeutics [118–120].

4.2. On–off delivery

Protein and peptide release can be engineered to occur in apulsatile mode, intended as the rapid and transient release of acertain amount of drug molecules within a short time-periodimmediately after a pre-determined off-release interval. Oneway to classify pulsatile DS is based on the physicochemicaland biological principles that trigger the release [121]. Thesedevices are classified into “programmed” and “triggered”DS. Inprogrammed-DS, the release is completely governed by theinner mechanism of the device (for example lag-time prior todrug release in some DS). In triggered-DS, the release isgoverned by changes in the physiologic environment of thedevice (i.e. self-regulated DS or biologically-triggered DDS) orby external stimuli (i.e. externally-triggered systems). In thelatter case, external stimuli, such as temperature changes,electric or magnetic fields, ultrasounds and irradiation, activatedrug release [122,123].

4.2.1. Programmed delivery systemsIn the case of programmed-DS, precisely timed drug delivery

can be accomplished by the spontaneous hydrolysis (i.e. bulkand surface eroding systems) or enzymatic degradation of thepolymer comprising the device.

Bulk- and surface-eroding systems may be engineered toachieve pulsed protein delivery slightly modifying the compo-sition of the device, which can be based on PLGA [124–126],cross-linked hydrogels [127], polyanhydride [128,129], and allthose biodegradable polymers discussed above. In case ofPLGA-based microparticles, more than an effective “pulsed”drug delivery, a booster release occurring over a period ofseveral weeks after a typical lag-phase has been realized[124,125,130,131]. A real “pulsed” protein delivery from PLA/PLGA-based devices has been achieved by Langer et al., whodeveloped a resorbable microchip based on PLA [132]. Inperspective, the implant can enable the patterned delivery ofmultiple agents.

Also surface eroding polymers, such as poly(anhydrides),can be of help when pointing to pulsed protein delivery [129].Recently, a polyanhydride-based laminated device has beenapplied to the pulsatile release of parathyroid hormone (PTH)[129]. The implantable DS, consisting of drug layers, isolation

layers and sealant filling, allowed multi-pulse release profiles ofPTH and bovine serum albumin (BSA) in their bioactive forms.This implant can be produced in various shapes and delivermore than one drug. Furthermore, the load of therapeutics canbe easily tailored over a broad range in the drug layers.

4.2.2. Triggered delivery systemsSelf-regulated DS (i.e. biologically-regulated DS) are

closed-loop controlled release devices in which the releaserates are adjusted by the system, in response to feedbackinformation, without any external intervention. This is the caseof pH-responsive systems, which have been mainly investigatedfor oral protein delivery [133,134].

The most interesting class of self-regulated DS for TEapplications is probably represented by biomolecule-sensitivehydrogels, a kind of biologically-inspired materials able toresponse to specific physiological stimuli, such as increase ofglucose levels or the presence of special proteins and/or enzymes[122,123]. These systems can be potentially manufactured inform of fibers, gel, sheets or microparticles to fabricate scaffolds.A great deal of interest has been focused to glucose-responsiveinsulin delivery since the development of pH-responsivepolymeric hydrogels that swell in response to glucose [135].The “intelligent” system consists of immobilized glucose oxidasein a pH-responsive polymeric hydrogel, enclosing a saturatedinsulin solution. As glucose diffuses into the hydrogel, glucoseoxidase catalyzes its conversion to gluconic acid, therebylowering the pH in the microenvironment of the membrane,causing swelling and insulin release. Recent progresses have beenmade in designing “smart” hydrogels able to specific recognize abiomolecule through molecular imprinting techniques [136].

Contrariwise to self-regulated DS, externally-regulated DSare open-loop controlled devices in which drug release can beactivated by an external stimuli, including temperature changes,magnetism, ultrasound, electrical effect and irradiation[122,137,138]. These systems make use of “smart” polymericmaterials, which respond with a considerable change in theirproperties to small changes in their environment. The incorpo-ration of these devices into polymer scaffold may potentiallyoffer the opportunity to control the time and space environmen-tal conditions by mean of an externally-imposed field.

5. The issue of delivery system integration in three-dimensionalscaffolds

A clinically effective TE approach requires a chrono-programmed scaffold, able to trigger the on-demand release ofmolecular agents fulfilling the specific needs of the bio-integrating tissue. This approach is part of the flourishingconcept of chrono-biotechnology which may be accomplishedthrough the tight control over dosing and localization ofsignaling molecule exposure in a complex, mostly anisotropic,dynamic three-dimensional environment [17,78].

To date, several attempts have been made to obtain systemsintegrating delivery devices and TE templates able to mimicECM and directionally reorganize tissue [139]. In the case ofmicrosphere-integrated scaffolds, however, a very limited

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number of the possible protein delivery strategies describedabove have been exploited. Some relevant results have beenachieved by the group of Mooney [17], who developed PLGAscaffolds for the sequential release of multiple GFs by mixingfree VEGF with empty and PDGF-loaded polymer particles andsubsequently assembling them into a porous scaffold [140].More recently, they have also presented an anisotropic systembased on a porous bi-layered PLGA scaffold able to expose onlyVEGF in one spatial region, and deliver VEGF and PDGF in anadjacent region [141]. In a similar attempt, PLA microparticlesplasticized with PEG were sintered into scaffolds formed byprotein-free and protein-loaded layers, thus allowing a releaseof different bioactive molecules restricted to specific regionswithin the scaffold [142]. These scaffolds may find utility inapplications where GF gradients or a region-dependent tissuegrowth are required.

Despite these encouraging results, important technologicallimitations exist. The major issue relies on the use of GF-loadedmicrospheres which are partly modified when formed intoscaffolds, thus altering their architecture and consequentlyrelease features. Furthermore, to engineer dynamic gradients ofa signaling molecule, the detailed understanding of releasekinetics at the single microsphere level is necessary.

An alternative approach to create microsphere-integratedscaffolds able to regulate both temporally and spatially GFrelease kinetics may take advantage of micromanipulation-basedtechniques. The simple dispersion ofmicrospheres within gel-likescaffolds is a well-established approach to achieve a temporalcontrol over GFs release [139]. It has been recently demonstratedthat through the fine tuning of microsphere formulation andscaffold properties it is possible to realize platforms able to controlthe microenvironmental conditions in terms of time evolution ofbioactive molecules delivery [143]. Possible developments ofthese findings may benefit from advances in micro- andnanotechnologies so as to engineer templates embedding micro-spheres releasing GFs at known release rates in a predeterminedand optimized spatial distribution within the scaffold. Actually,devices acting as single point source may be micropositioned by3D printing and soft lithography to obtain highly regulatedstructures able to trigger the extent, and possibly the architecture/structure of tissue formation [10,144]. The combination ofmicropositioning systems and mathematical modeling describingthe complex and multiple mechanisms governing the releasekinetics from single microspheres within the scaffold can be ofhelp to realize scaffolds with highly controlled architecture bycomputer-aided scaffold design (CASD) [10,145].

A possible limitation of DS-integrated scaffolds derives fromtheir pre-defined nature. In fact, once pre-programmed in vitro,they will not be able to interactively modify release kineticsaccording to the needs of the surrounding tissue. As underlinedabove, a more effective biomimicry could be obtained if a dual-reciprocity scheme could be encoded in the matrix. In this way,cells can trigger the on-demand development of ECM and, inturn, the engineered scaffold could stimulate cell behaviorsthrough the controlled release of bioactive molecules. Inperspective, the use of bottom-up strategies based on molecularself-assembly appears very promising. On this matter, the group

of Stupp has developed a class of peptide amphiphile (PA)molecules that self-assemble into three-dimensional nanofibernetworks under physiological conditions in the presence ofpolyvalent metal ions [37]. While PA self-assembly entrapscells in the nanofibrillar matrix, the entrapped cells internalizethe nanofibers and possibly utilize PA molecules in theirmetabolic pathways. The method is not limited to uniaxialalignment but can be used to guide self-assembling nanofibersaround corners and in complex patterns. It is also versatileenough to be used in the alignment of other self-assemblingsupramolecular systems starting from solutions of smallmolecules [146].

It should also be mentioned that a multifunctional scaffoldshould not only provide a controlled administration of relevantbiomacromolecules and their gradients, but also present suchmolecules in a suitable conformation state, mimicking ECM-GFs binding. Indeed, it has been shown that molecularlydecorated materials enhance tissue formation through themodulation of the interaction between protein signaling andbiomaterials appears to be fundamental to provide a betterintegration of the scaffold with the neo-forming tissue [147].

These emerging approaches suggest that the next-futurescaffolds will not be realized by simply integrating DS withinthe scaffolds. Indeed, taking advantage of the currentknowledge of drug delivery and biomaterial science, multi-functional scaffolds, where the polymer three-dimensionaltemplate itself acts as a biomimetic, programmable and multi-drug delivery device, should be designed.

6. Conclusions

Extraordinary progresses have been made in the last decadetowards the design of scaffolds with a suitable multiscalehierarchical structure and the design of DS able to release activeproteins according to virtually any complex delivery pattern.The integration of cutting-edge scaffold production technolo-gies and DS may lead to significant advances in boththerapeutic applications of TE and basic knowledge on cellguidance and tissue morphogenesis. However, technologicaland scientific challenges have still to be overcome to realize afaithful mimic of the complex orchestration of structural andmolecular signals presented by the natural ECM. For instance,micro- and nano-technologies should be further exploited inorder to organize a DS-integrated biomaterial scaffold with therequired spatial resolution. On the other hand, design of futurebiomaterial platforms should benefit from a quantitativeunderstanding of the influence of the amounts of morphogensand their gradients on cell fate, as well as the influence ofdifferent microenvironments on their action.

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