Macroscale delivery systems for molecular and cellular payloads.

| FOCUS | INSIGHT ARTICLETYPE REVIEW ARTICLES PUBLISHED ONLINE: 23 XXOCTOBER MONTH YEAR | DOI: 2013 | DOI: 10.1038/NMATXXXX 10.1038/NMAT3758

Macroscale delivery systems for molecular and cellular payloads Cathal J. Kearney and David J. Mooney* Macroscale drug delivery (MDD) devices are engineered to exert spatiotemporal control over the presentation of a wide range of bioactive agents, including small molecules, proteins and cells. In contrast to systemically delivered drugs, MDD systems act as a depot of drug localized to the treatment site, which can increase drug effectiveness while reducing side effects and confer protection to labile drugs. In this Review, we highlight the key advantages of MDD systems, describe their mechanisms of spatiotemporal control and provide guidelines for the selection of carrier materials. We also discuss the combination of MDD technologies with classic medical devices to create multifunctional MDD devices that improve integration with host tissue, and the use of MDD technology in tissue-engineering strategies to direct cell behaviour. As our ever-expanding knowledge of human biology and disease provides new therapeutic targets that require precise control over their application, the importance of MDD devices in medicine is expected to increase.

M

acroscale drug-delivery devices are modernizing the ancient Ars Medicine (‘art of healing’) by using engineering principles to exert spatiotemporal control over drug availability. Today, most medications are administered orally or by intravascular/intramuscular injections, and the circulatory system transports them throughout the body. Macroscale drug-delivery devices instead deliver a bioactive agent at the desired treatment site by means of a carrier material whose physical and chemical properties control the presentation of the agent. Macroscale drug-delivery devices are distinguished from particulate drug-delivery devices, such as micro- and nanocarriers, by their size (at least one dimension over 1 mm). Different types of material (for example polymers) in varying physical forms, such as gels, are used to fabricate MDD devices that are delivered by implantation or injection. The drug(s) or bioactive agents delivered include small-molecule drugs, proteins, oligonucleotides, silencing RNA, plasmid DNA and antibodies. Within this MDD classification, we also include devices that maintain control over the presentation of contained microcarriers, nanocarriers and cellular payloads. For example, in certain gene-therapy applications these microcarriers may be genetically modified cells that act as a source of various factors1. An early observation in the 1960s laid the foundation for MDD technologies, when it was demonstrated that the diffusion rate from a drug reservoir could be controlled using a silicon rubber membrane2. This became the basis for the development of several drug-delivery products in applications such as glaucoma, intrauterine devices (IUDs) for contraception, and skin patches that could emit drugs at constant rates3. These devices originally focused on the delivery of small-molecule drugs, but as knowledge of the role of growth factors and other macromolecules in localized biological signalling increased, this motivated the delivery of sensitive biological agents directly to a repair site. In the mid-1970s it was demonstrated that (large) protein drugs could be gradually released from macroscopic carriers, leading to the development of synthetic polymer systems to control the spatiotemporal presentation of proteins while protecting them before their release4. These early findings formed the basis for the field of macroscale drug delivery, which has grown in parallel with advances in tissue engineering, stem cell biology and our knowledge of biological pathways. We begin by highlighting the key advantages of MDD systems. The principal mechanisms that control agent

release or presentation — namely diffusion, affinity and material degradation — are introduced, and salient examples are provided. We then describe mechanisms that allow one to trigger release externally, either by affecting these various mechanisms or by stimulating local convection, as well as systems that immobilize factors. We provide general guidelines that help to direct the choice of material and control mechanism depending on the application and delivered agent. We also describe multifunctional devices that exploit MDD techniques to provide solutions to complex clinical problems. Finally, we discuss current challenges facing the field.

Why use macroscale biomaterials?

When drugs are delivered without control over the location or rate of delivery, large doses are frequently required to stimulate a therapeutic effect. These high doses, beyond the waste and added expense, can lead to increased toxicity or undesirable side effects. Drug-delivery devices — both macroscale and particulate — overcome the limitations of classic drug administration in three ways (Fig. 1). First, they are either delivered locally or have targeting mechanisms that provide strategic and precise spatial control. Once in place, the devices act as a depot of the bioactive agent(s) at that site. This localization greatly reduces the required dose, and often the toxicity. Second, they provide temporal control over the presentation of the bioactive agents, sustaining therapeutic concentrations over a longer time period than if the factors were administered without a carrier, or provide more sophisticated patterns of presentation (for example sequences of drugs or pulsatile presentation). Finally, drug-delivery systems can protect sensitive biological agents, shielding them from degradation until they are released. In fact, the short half-lives of many modern therapeutic candidates emphasize the importance of this feature. Macroscale drug-delivery devices can circumvent challenges associated with systemic particulate delivery: namely, renal clearance, poor targeting efficiency, and degradation by serum nucleases. Further, macroscale systems allow one to make effective use of controlled release mechanisms that have length-scale dependencies (for example diffusion). Finally, multifunctional MDD technologies result from integrating MDD into existing medical devices or using them to regenerate tissues. In contrast, the facile intravenous injectability of many particulate systems makes them attractive in a variety of clinical applications where the advantages of MDD systems are not

School of Engineering and Applied Sciences, Harvard University, Cambridge, Massachusetts 02138, USA, and Wyss Institute for Biologically Inspired Engineering at Harvard University, Boston, Massachusetts 02115, USA. *e-mail: [email protected] 1004

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required. Nano- and microcarrier systems that aim to circumvent the challenges described above are reviewed elsewhere5–7.

Mechanisms of spatiotemporal control

Target site

Macroscale drug-delivery materials can generally be placed into one of three mechanistic categories (Fig. 2), depending on whether release is controlled by diffusion, drug–carrier affinity or degradation of the material. In practice, the release profile is often a combination of these, with one mechanism being dominant. Convection provides an additional mechanism, accessed by applying external stimuli that mechanically actuate materials, and this can increase and/or direct drug release. The contribution of each of these mechanisms can be interpreted using the general mass-balance equation (Fig. 3). An alternative approach is to immobilize factors onto the MDD material. This provides precise and stable control over the spatial pattern of drug presentation and can maintain the activating signal for extended periods. In what follows we provide a brief description of each mechanism as well as relevant examples (see also Table 1). Diffusion-controlled release. Diffusion controls the release of a drug when diffusion in the carrier occurs slowly compared with the rate of drug dissociation from the material, yet it occurs much faster than material degradation. In diffusion-based systems, drugs are typically encapsulated into an internal reservoir or mixed homogeneously in the material. The earliest example of a MDD device was a porous silicon membrane through which anaesthetic gases encapsulated in a reservoir slowly diffused2,3. Drug release from these systems can be represented using simple mathematical models, predicted and manipulated. This mode of delivery has been exploited in several devices, including Ocusert (controlled release of pilocarpine for glaucoma), Progestasert (contraceptive IUD) and transdermal delivery patches used in applications such as nicotine administration (for instance, Nicoderm CQ) or pain relief (for example, Duragesic). The release rates from diffusion or reservoir devices are typically independent of carrier concentration, as the drug inside the reservoir is loaded at a concentration much greater than the solubility in the polymer membrane, and the membrane can be engineered to control the release rate by adjusting the drug solubility in the polymer, the drug diffusivity through the polymer and the partition coefficient (ratio of drug solubility in membrane to drug solubility in surrounding media)8,9. Some devices, particularly those used in transdermal delivery, have semipermeable channels that exert control over larger drugs. For these, release is controlled by the ratio of pore size to the drug’s molecular volume, as well as by the tortuosity of the pores10. Physical characteristics of the device can also be adjusted to control the release rate. For example, in some of the earliest papers on this topic it was demonstrated that increasing the membrane thickness decreased the release rate2, and that altering the surface area of the device adjusted the rate10. Fabrication of these devices is generally straightforward: once the polymer membrane, thickness and shape have been chosen, the reservoir is loaded with drug (often in a liquid or gel carrier to ensure good interfacial contact with the polymer) and the device is sealed by heat treatment or an adhesive. (See Box 1 for a discussion of appropriate choice of material and control mechanism for a particular drug.) Matrix systems are often used instead of reservoir systems when significant danger is associated with accidental rupture of the membrane and subsequent rapid drug release. Hydrogels, which consist of hydrated networks of crosslinked polymers, have been used extensively to fabricate matrix systems. Numerous natural materials, including alginate11, gelatin12, chitosan13 and fibrin14, and synthetics such as poly(ethylene glycol) (PEG), poly(hydroxyethyl methacrylate) (PHEMA) and poly(vinyl alcohol) (PVA) have been used in matrix systems15. In these MDD systems, the diffusion

MDD at target site provides: • Localized signal • Controlled/retarded release • Protection for labile drugs

Enteral/parenteral administration: • Typically reqires larger doses, increasing toxic side effects • Can expose drugs to harsh environment • Often requires repeated doing

Figure 1 | Enteral/parenteral administration compared with MDD. Drugs delivered orally (enteral administration) are exposed to various acidities as they pass through the gastrointestinal tract before being absorbed into the bloodstream. Parenteral delivery (injection into the bloodstream, the peritoneal cavity, subcutaneously, or intramuscularly) often requires the drug to enter and survive the systemic circulation and to avoid renal clearance, with the result being that only a small percentage of the drug typically reaches the target site. Consequently, large and repeated dosing is often necessary, which can increase the risk of toxicity. Macroscale drugdelivery technologies allow a carrier to provide a drug reservoir at a desired anatomic site (typically the wound site), with a continuous or chosen temporal pattern of drug presentation. Macroscale drug-delivery can eliminate or reduce the requirement for repeat dosing, while protecting the drug prior to its release to allow long-term dosing (body outline adapted with permission © NIDDK, NIH).

coefficient of the drug through the matrix is a function of steric interactions between the drug and the matrix. When the pore size is much larger than the size of the drug, diffusion is a function of the tortuosity of the release path. Instead, when the pore size is comparable to the size of the drug, steric hindrance from the polymeric chains reduces the diffusion coefficient. These parameters can be used to guide the selection of the matrix for a specified drug: tortuosity can be increased and/or pore size decreased by increasing the polymer concentration (as studied in alginate beads16) or the molecular weight of the chains17, and by increasing crosslinking (as studied in poly(acrylic acid)18,19). Because of the versatility of these systems, they have been used to deliver therapeutic candidates that cover multiple length scales, ranging from classic small-molecule drugs, to DNA and protein therapeutics20–24. And because of the complexity of their release process, there has been extensive modelling and computational work in an effort to predict the kinetics of drug release from a given system, as recently reviewed17. Fabrication of these systems is achieved by physical or chemical crosslinking. Physical crosslinking can be attained by exploiting a number of features, including hydrophobic interactions at body temperature for amphiphilic polymers (such as poly(ethylene oxide)–poly(propylene oxide)–poly(ethylene oxide) and PEG– poly(lactide-co-glycolide)–PEG), charge interactions (such as in ionically crosslinked alginate, chitosan and glycerophosphate), hydrogen bonding (for instance in gelatin/agar and hyaluronic acid/ methylcellulose) and stereocomplexation (as in poly(l-lactide) and poly(d-lactide))25. The advantage of physically crosslinked polymers is their ability to form in situ. Some of these polymers, however, are



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b

Nicoderm CQ

NATURE MATERIALS DOI: 10.1038/NMAT3758 Fibrin

Fibrinogen

FpB FpA FpB FpA

Diffusion control

Fibres

Nicotine-containing layer

γ-module β-module b a Coiled coil

γ-module β-module a b

Backing layer

Coiled coil

Thrombin cleaves FpA and FpB

Rate-controlling membrane (HDPE) Adhesive Skin

D-region b a

10 µm

E-region

Diffusion-controlling fibre mesh

D-region D-region b b a a

a b D-region

E-region

D-region b a

b a D-region

E-region

Protofibril

c

d Layer-by-layer coating

Poly(lactic-co-glycolic) scaffold

Degradation control

Glycolic acid O O HO

H O

X

Y

O

Scaffold

HOOC O HO

N

N

OR1

OR2

O

O O OR3

O

O

H

NH CH3

OH

n

Ester linkages undergo hydrolysis

n

Chondroitin sulphate: can undergo enzymatic breakdown

f β-Cyclodextrin coating

Alginate gel –OOC

–OOC

OH

O

2 cm

O

LbL coating

Ester linkages undergo hydrolysis

1 mm

Poly(β-aminoester) O

H

O

e Affinity control

(+) (–) Drug (+) (–) (+)

Lactic acid

O

OH

O

O

OH

OH G

O

O OH

HO

OH O

OH O

–OOC

Hexamethylene diisocyanate

O

N

H

H

O O O C

–OOC G

O C N H N

M

M

100 µm

5 mm Ca2+ cooperatively bind to crosslink polymer chains

O

O

Hydrogen bonding

N CH O O N H

N H

O C

O

O O C O

N O H C N O O H C

Cyclodextrin polymer with hydrophobic pocket acts as affinity site

Van der Waals interaction

Figure 2 | Devices exemplifying the different release mechanisms. a,b, Diffusion control. c,d, Degradation control. e,f, Affinity control. a, Nicoderm CQ, a patch that helps smokers quit, consists of a reservoir of nicotine that diffuses through a rate-controlling membrane to be absorbed into the body through the skin. b, Scanning electron microscopy (SEM) image of fibrin gel used in matrix systems. The fibrinopeptides A and B (FpA and FpB) are cleaved by thrombin, and the fibrin monomers form protofibrils; lateral interactions with other protofibrils result in fibres that mesh to form a network to retard diffusion. c, Photomicrograph of a PLG carrier that can undergo hydrolytic degradation and release contained factors (inset: SEM image of the device). Ester linkages between the lactic and glycolic acid groups undergo hydrolysis, and the drug is released as the device erodes. d, Layer-by-layer (LbL) coating of chondroitin sulphate/poly(β-aminoester)-2 on Therics 3D-printed scaffolds consisting of polycaprolactone/b-tricalcium phosphate (PCL/BTCP) copolymer blend. The coating released rhBMP-2 following degradation of the ester linkages in the poly(β-aminoester). e, Alginate gel with mitoxantrone subcutaneously injected above the hindlimb (skin removed to expose site), and the gel being ejected from a syringe (inset). Ionically crosslinked alginate gels are formed with divalent cations that bind sites on the G blocks of two adjacent chains. Cationic drugs and heparin-binding drugs both show affinity for alginate chains, which can be exploited to control drug release. f, SEM of β-cyclodextrin coating on a polyester mesh (Parietex), which is used to control release of hydrophobic antibiotics. The cyclodextrin hydrophobic pockets can be incorporated into a hexamethylene diisocyanate polymer network and used to coat devices. Figure reproduced with permission from: b, ref. 14, © 2009 RSC; d, ref. 46, © 2011 Elsevier; f, ref. 33, © 2010 Elsevier.

mechanically weak and become unstable after several hours in vivo. Alternatively, polymers can be covalently crosslinked by small molecules (for example, genipin has been used to crosslink amino-terminated PEG, gelatin and albumin) or other polymers (for instance hydroxyethyl methacrylate (HEMA) crosslinked with ethylene glycol dimethylacrylate (EGDMA)). Radical chain polymerization can also be used to form polymer matrices. For example, photocrosslinking of 2-hydroxyethyl methacrylate and diethylene glycol dimethacrylate (DEGDMA) using ultraviolet light and photoinitiators (such as Irgacure 651) has been used to form drug-loaded, multilaminated, diffusion-controlled release systems26. The drug can be mixed with the primary polymer prior to crosslinking, or, if the crosslinking technique requires an environment that may have adverse effects on the drug, it can be soaked into the hydrogel after formation. 1006

Drug-release systems can be designed to respond to local environmental cues (such as pH, temperature or enzymes) that act on the carrier to increase drug diffusion3,7,15,27,28. For example, an interpenetrating network of temperature-sensitive poly(N-isopropylacrylamide), which undergoes a phase transition and rapid deswelling at 32 °C, and the pH-sensitive poly(methacrylic acid) (PMAA), which swells at pH > 5.5, demonstrated pH- and temperature-dependent diffusional control over release of model drugs of different molecular weights29. Approved devices based on this concept include Eudragit, a carrier that becomes permeable at select pH values and consequently releases drug at a specific location along the gastrointestinal tract. Affinity-controlled release. When a drug has a strong, reversible chemical interaction with the material used to fabricate the MDD NATURE MATERIALS | VOL 12 | NOVEMBER 2013 | www.nature.com/naturematerials




Mass balance/diffusion equation 2C – Ci v + Ri i Diffusion Convection Chemical reactions

Diffusion control

∂Ci ~D ∂t

b

Δ

a

Δ

∂Ci = D ∂t

Δ

system, this affinity can control drug release. These interactions are usually based on ionic bonds, hydrogen bonds, van der Waals forces or hydrophobic interactions30. Affinity-based interactions may result from simple mixing of the drug and polymer during preparation31,32, or by adding the drug to the polymer carrier following synthesis33. Heparin sulphate is a highly sulphated anionic polysaccharide found in the extracellular matrix (ECM) that plays a role in immobilizing and releasing proteins involved in many cellular processes34. Heparin — its therapeutic derivative that shares many of the same properties — has been used for affinity-controlled release of heparin-binding proteins (for instance vascular endothelial growth factor (VEGF) or basic fibroblast growth factor (bFGF)) from heparin-containing gels (such as hyaluranon–PEG–heparin gels)32. Heparin was incorporated into these gels by adding thiol groups on the hyaluranon and heparin, and subsequently co-crosslinking them with PEG diacrylate32. Alginate is a natural anionic polysaccharide that has also demonstrated affinity-controlled release of many heparin-binding proteins, including VEGF and bFGF35. Delivery of angiogenic factors with both of these hydrogel systems led to a pro-angiogenic response in murine models of peripheral artery disease, with rescue of hind limbs that would otherwise necrose32,35. The affinity of alginate hydrogels for heparin-binding proteins can be further increased by sulphonation of the uronic acid monomers. In fact, sulphated alginate gels resulted in slower release of bFGF from gels36. Furthermore, cyclodextrins — cyclic oligosaccharides with a hydrophobic core and hydrophilic exterior — have been studied for affinity-controlled release of hydrophobic small molecules30. A significant extension in release time of hydrophobic antibiotics (for example, Rifampin, Novobiocin and Vancomycin) has been reported from β-cyclodextrin as compared with a dextran-based system33. Beta-cyclodextrin gels have been formed by direct crosslinking using isocyanate crosslinkers, and hydroxypropyl-β-cyclodextran has been crosslinked with citric acid using an esterification reaction30. From an engineering standpoint, because the control comes from interactions between the drug and the matrix, this is the target for adjusting release rate. Increasing or decreasing the affinity between both molecules leads to a concomitant increase or decrease in release rate. For example, the cationic chlorpheniramine maleate had a slower release rate than the anionic sodium salicylate from the polyanionic alginate37. Additionally, as the number of interaction sites in the carrier increases, the probability of a drug exiting to the surface decreases. Therefore, either decreasing the number of sites of interaction or increasing the drug concentration will result in a higher release rate. In recent years, this concept of adjusting the affinity has been explored further with molecular imprinting (MIP) — an approach that creates templated cavities, within the carrier material during device fabrication, to match specific drug molecules30. The earliest example of MIP used methacrylic acid (MMA) as the functional monomer (the carboxylic acid functional groups form ionic interactions with amines and hydrogen bonds with polar groups) and EGDMA as the crosslinker to imprint theophylline (bronchodilator used in asthma) and diazepam (tranquillizer used in anxiety-related disorders)38. Methacrylic acid crosslinked with EGDMA has also been used in combination with hydrogel materials to form soft contact lenses that control the delivery of timolol for glaucoma treatment over a 12-h period. These lenses could be reloaded in a timolol solution overnight for reuse the next day39. Protein delivery is more challenging from MIP systems owing to steric hindrance and the large and flexible protein template that results in poorly defined binding sites40. As an alternative, select peptide fragments of the full protein can be used as the template in porous carriers that allow diffusion of large proteins. The resulting MIP system can then bind either the peptide or the full protein in an ‘epitope’ approach, as demonstrated with release of oxytocin (a neurohypophyseal hormone) from MMA-based macroporous carriers41.

2C i

Degradation control

∂Ci ~ Ri ∂t

ace urf

S

Δt Bul

ke

c

n

sio

ero

ros i

on

Affinity control

∂Ci ~ Ri ∂t

d

Immobilized factors

∂Ci = 0, ∂t Ci = Constant

Bioactive agent

Reversible bond

Crosslinked hydrogel

Permanent bond

Figure 3 | Control exerted by drug-delivery devices can be mathematically modelled using the generalized mass-balance equation. This equation represents how the concentration of a drug changes over time as a function of its diffusion, convection and chemical reactions. In the equation, Ci is the concentration of drug i, t is time, D is the diffusivity of the drug through the carrier, v is the volume-averaged velocity, and Ri accounts for chemical reactions (such as drug/carrier degradation, or drug-vehicle affinity interactions). In the absence of flow, which is typically assumed for most systems, the convection term is zero and the diffusion and/or chemicalreaction term controls the drug-concentration kinetics. a, Diffusion control occurs if the timescale of diffusion is longer than that of drug-carrier dissociation, yet shorter than that of material degradation. Steric interactions and the tortuosity of the release path control the release kinetics. b, In degradation-controlled systems, the chemical-reaction term dominates, with drug being released following carrier breakdown. Degradation can occur primarily with carrier mass loss moving from the outside in (surface erosion) or evenly throughout the entire carrier (bulk erosion). c, Similarly, affinitycontrolled systems are rate-limited by the chemical reaction mediating the reversible interaction between drug and carrier. Drug-binding retards the drug from freely diffusing from the carrier. d, In systems with immobilized factors, all terms on the right side of the equation are equal to zero, as the drug is bound to the carrier and the concentration is constant over time.



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Table 1 | Selection of clinically used MDDs and their control mechanism. Product name

Therapeutic application

MDD rate-controlling material

Control mechanism

Drug delivered

INFUSE

Spinal fusion; tibial-fracture treatment; sinus augmentation; vertical ridge augmentation

Collagen type I

Affinity control

Recombinant human bone morphogenetic protein-2 (rhMP-2)

OP-1 Putty/Implant

Spinal fusion; tibialfracture treatment

Collagen type I

Affinity control

Recombinant human bone morphogenetic protein-7 (rhMP-7)

Gliadel

Surgery adjunct for malignant glioma

Poly(carboxyphenoxy propane: sebacic acid), (20%:80%)

Degradation

Carmustine

Zoladex

Prostate cancer, endometriosis and breast cancer

Polylactide/poly(lactide-co-glycolide)

Degradation

Goserelin acetate

Zilver PTX

Peripheral artery reperfusion and anti-restenosis

Nitinol stent (polymer-free)

Matrix diffusion

Paclitaxel

Medusa

Hepatitis C treatment

Polyglutamate/vitamin E

Matrix diffusion

Interferon-α

AzaSite

Bacterial conjunctivitis

Poly(acrylic acid)

Matrix diffusion

Azithromycin

Besivance

Bacterial conjunctivitis

Poly(acrylic acid)

Matrix diffusion

Besifloxacin

Taxus

Coronary artery reperfusion and anti-restenosis

Poly(styrene-b-isobutylene-b-styrene)

Matrix diffusion

Paclitaxel

Cervidil

Cervical ripening for induction of labour

Poly(ethylene oxide) and urethane

Matrix diffusion

Dinoprostone (synthetic prostaglandin)

Endeavour


Phosphorylcholine

Matrix diffusion

Zotarolimus

Xience V


Fluoropolymer

Matrix diffusion

Everolimus

CYPHER


Poly(ethylene-co-vinyl acetate) and poly(n-butyl methacrylate) (PBMA); outer PBMA coating also controls diffusion

Matrix- and reservoir-diffusion controlled

Sirolimus

ATRIDOX

Chronic periodontitis

Poly(d-l-lactide)

Matrix diffusion/ degradation

Doxycycline hyclate

Absorb


Poly(d,l-lactide)

Matrix diffusion/ degradation

Everolimus

Ocusert

Glaucoma treatment

Poly(ethylene vinyl acetate)

Reservoir diffusion

Pilocarpine

Progestasert

Intrauterine contraceptive device

Poly(ethylene vinyl acetate)

Reservoir diffusion

Progesterone

Nicoderm CQ

Smoking cessation

High-density polyethylene

Reservoir diffusion

Nicotine

Duragesic

Opioid for pain medication

Ethylene-vinyl acetate copolymer

Reservoir diffusion

Fentanyl

Degradation-controlled release. If drug diffusion or dissociation of the drug is slow compared with the degradation of the carrier material, degradation may be the primary mechanism controlling drug release. Degradation may occur preferentially on the surface of the carrier or simultaneously throughout the carrier body. Mechanisms include dissolution, hydrolytic cleavage of the polymer and enzymatic degradation15. Typically, polymer degradation results in erosion — that is, material loss from the polymer bulk. Many different polymers with carefully selected degradation rates and physical properties that match the application period (for example poly(lactide-co-glycolide) (PLG), poly(glutamic acid) and collagen) have been used to form devices. Degradation-controlled release has been exploited for several FDA- (Food and Drug Administration) approved oral medications, such as those based on SkyePharma’s Geomatrix technology. Degradation-regulated release has also been exploited in a number of preclinical studies. For example, sequential delivery of VEGF followed by plateletderived growth factor (PDGF) from PLG in a murine model of peripheral artery disease was used to promote angiogenesis and vessel maturation42. The release kinetics were controlled by either the processing technique, or by selecting polymers that degraded at different rates. Using PLG as an illustrative example, increasing the molecular weight of the polymer results in slower erosion, as does 1008

increasing the crystallinity or ratio of lactide:glycolide. Lactide is more hydrophobic than glycolide and slows hydrolytic breakdown, whereas increasing crystallinity slows water diffusion into the bulk. The more rapidly degrading polymer generally results in faster drug release. Enzymatic degradation of chitosan has also been used to control release of heparin-bound angiogenic factors (fibroblast growth factor-1 (FGF-1) and FGF-2) contained in a gel in order to stimulate angiogenesis43. Multiple layers of different carrier–drug combinations can be used to achieve sequential release of factors in degradation-controlled systems44–46. Carrier materials that show altered degradation in response to environmental cues can also be used to manipulate drug-release rates. For example, polyanhydrides exhibit pH-dependent degradation, and this was used to control release of p-nitroaniline from poly(bis(p-carboxyphenoxy)propane anhydride)47. Externally triggered release. In numerous pathological conditions, the required dosing can be difficult to predict a priori, and may need to be actively managed. Real-time control may allow one to ‘dial in’ the dose needed at each instant in time. Macroscale drugdelivery systems have been developed that respond to a variety of external cues, including magnetic and electric fields, ultrasound and NATURE MATERIALS | VOL 12 | NOVEMBER 2013 | www.nature.com/naturematerials


NATURE MATERIALS DOI: 10.1038/NMAT3758 wireless signals48–51 (Fig. 4a). In ultrasound-responsive materials, the ultrasound signal is typically used to accelerate the degradation of the MDD carrier by producing cavitation, thereby accelerating release during the ‘on’ phase. When the stimulus is removed, release rates can remain elevated or return to baseline levels49,52. In one of the earliest demonstrations of ultrasound-stimulated drug release, the release of p-nitroaniline and bovine serum albumin from polylactide, polyglycolide and poly(bis(p-carboxyphenoxy)methane) (PCPM) was demonstrated in vitro49. Release of p-amino hippuric acid (PAH) from copolymers of bis(p-carboxyphenoxy)propane with sebacic acid (PCPP/SA) demonstrated the in vivo applicability of this approach49. External cues can also trigger convection if they result in gross and rapid structural changes to the carrier (Fig. 4b). In magnetically responsive systems, iron oxide particles that are dispersed throughout a hydrogel network deform the network in a magnetic field. Inclusion of micropores in the system results in large deformations, which lead to more exaggerated and rapid carrier collapse53. This results in convection within the device and accelerated release of incorporated agents, which can include chemotherapeutic drugs, plasmid DNA, proteins and stem cells53. Similarly, electroresponsive gels — typically, polyelectrolyte hydrogels that undergo deswelling

INSIGHT | REVIEW ARTICLES or bending in response to an electric field — can be used to control drug release54. Forced convection has led to the release of pilocarpine hydrochloride and raffinose from PMAA, with release following the temporal profile of deswelling, not the application of the field55. An alternative approach uses an implanted microchip device that internally processes information from an external trigger. Inclusion of many individually addressable depots within the device can allow for digital control over release of waves of the same drug or of multiple drugs from the same device. The prototype device in this category is a wireless, multireservoir microchip system that triggers electrochemical dissolution or electrothermal ablation of membranes that conceal drug reservoirs in response to input51,56. This system has recently undergone a successful first clinical trial, delivering the anti-osteoporotic human parathyroid hormone fragment (1–34) (hPTH(1–34)) to postmenopausal women for 3 weeks56. Immobilized factors. In some situations, one may want to provide a highly localized signal to cells that come in contact with the MDD device (for example to drive stem cell differentiation57), with little to no release of the signal to other cells in the environment. Signalling in this scenario occurs by direct interaction between cell and bound factor, without internalization or removal of the signalling molecule

Box 1 | Guidelines for selecting carrier materials.

The choice of a specific material and control mechanism for a particular drug has become blurred as researchers have often adapted their preferred carrier material(s) to exert control over multiple drug types. But there are still some general guidelines that can be used when designing and selecting materials for a specific application. Once the therapeutic application has been clearly defined, the key parameters in selecting the carrier material include the chemistry and physical properties of the drug, the target site, the target dose and the duration/temporal profile of drug release that is desired. Drug When the drug or drugs have been chosen, their chemical, physical and biological properties should be understood and integrated into the choice of material. The size of the drug can determine its suitability for diffusion-controlled release systems. For example, large molecules often have sufficient steric interactions to exploit in matrix-diffusion-controlled systems. Alternatively, small molecules may be more amenable to a reservoir system if they have suitable diffusivity and solubility in the controlling polymer. Certain physicochemical properties of the drug (for example charge, heparin-binding, hydrophobicity) are necessary for affinity-controlled release systems to be used (such as cyclodextrin systems for hydrophobic drugs or polyionic hydrogels for charged drugs). The materials challenge in this case is matching these drug properties to complementary carrier materials. Larger drugs (such as protein and DNA) that have charged and/or heparin-binding properties that can interact with a carrier matrix (such as alginate, chitosan or hyaluranon–PEG–heparin gels) are often best suited to matrixdiffusion or affinity-controlled release systems. Also, these systems typically help to protect labile drugs. Degradation-controlled systems also can confer protection for labile drugs and can be used for many drug types, assuming that the diffusion through the carrier can be sufficiently limited. Target site Both the accessibility of the selected therapeutic site to surgical implant or injection/catheter delivery and the need to retrieve the MDD following its lifetime have a large impact on the carrier

material selected. For example, early clinical diffusion-controlled MDD devices such as Ocusert (controlled release of pilocarpine for glaucoma), Progestasert® (contraceptive IUD) and transdermal drug-delivery patches do not require surgical implantation or injection, and can be easily removed after use. But because these devices are often not designed to degrade, implanting them surgically for other applications may require their removal after treatment. For this reason, degradation-controlled systems, or systems that control drug release by a different mechanism but ultimately degrade, may be more suitable. Matrix-diffusion and affinity-controlled systems are often amenable to injection, which makes them desirable from a delivery standpoint. For example, alginate163 and chitosan164 can be adapted to an injectable formulation, and collagen can be used as a thermosetting (at 37 °C) injectable polymer. Target dose, duration and temporal profile desired The required duration of drug availability and its temporal profile (such as continuous versus pulsatile) also affect the choice of material for controlled delivery. The duration of sustained release can be extended by exploiting the interaction between the drug and the carrier material. For example, in diffusion-controlled systems one can increase retention and slow release by reducing diffusivity and solubility. With affinity-controlled systems, the release can be slowed down by increasing the binding of the drug to the carrier material, or the ratio of binding sites to drug molecules. The duration of release in degradation-controlled systems can be extended by slowing carrier degradation. At the extreme end of this spectrum are the immobilized MDD systems, which are suitable for very-long-term to semi-permanent cell–factor interaction. If a system that is responsive to external stimuli is central to the therapeutic success, externally triggered MDDs that respond to ultrasound, magnetic or electric signalling can be designed. Finally, we note that the therapeutic potency of the loaded drug may drive selection: affinity-, diffusion- and degradation-controlled systems are typically limited to delivering mass quantities smaller than that of the carrier polymer (typically from nanograms to micrograms). Instead, reservoir systems can be used to store larger quantities of drug.



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Δ

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Open pores Bioactive agent Crosslinked hydrogel

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Figure 4 | Externally triggered MDD systems control the instantaneous release rate in response to energy transmitted into the system. a, External energy sources (such as ultrasound) can accelerate drug release during the ‘on’ phase by altering the microstructure of the material and accelerating degradation, accelerating drug–carrier dissociation, and probably introducing convection. b, In externally triggered systems that undergo rapid macroscopic deformation, convective flow is introduced and drug dissociation accelerate. SEM images, 2 mm wide. SEM images adapted with permission from ref. 53, © 2011 NAS.

from the surface. This can maintain bioactivity for extended periods, and also allows for co-stimulatory signalling and clustering. A key challenge to this approach is to maintain the functionality of the bound factor on immobilization, and this requires careful consideration of the coupling chemistry. One approach uses a tethering chain that covalently binds specific functional groups on the factor (for instance amine group in epidermal growth factor58) and on the carrier material. For example, transforming growth factor-β (TGF-β) was incorporated into PEG hydrogels by covalently bonding acryloyl-PEG-N-hydroxysuccinimide to primary amines on the TGF-β. Smooth muscle cells cultured on these carriers secreted more ECM proteins in response to the tethered TGF-β than to the equivalent concentration of soluble factor59. Similarly, a PEG spacer was used to tether bone morphogenetic protein-2 (BMP-2) to macroporous PLG carriers, and this significantly outperformed controls when implanted in full-thickness cranial defects in rabbits60. A second approach incorporated binding sites in peptide sequences61 — or, using genetic engineering, in growth factors62,63 — that can be covalently and stably conjugated to carriers. Beta-nerve growth factor (β-NGF) was genetically modified to have a transglutaminase substrate of α2-plasmin inhibitor, which can be covalently crosslinked to fibrin by factor XIIIa (ref. 64). The crosslinking 1010

NATURE MATERIALS DOI: 10.1038/NMAT3758 domain was linked to the β-NGF via a plasmin-cleavable linker that can be enzymatically degraded by cells64. In an in vitro model, dorsal root ganglia grown in fibrin matrices demonstrated their ability to cleave the plasmin linker, and the released β-NGF enhanced neurite growth64. A similar system was tested in vivo, this time with BMP-2 as the tethered factor, and showed efficacy in a preclinical rat craniotomy model65.

Medical-device integration

Macroscale drug-delivery concepts are being integrated with existing medical devices to allow release of bioactive agents that work in concert with the device to provide improved and/or new functionality. This approach is being applied to multiple disease states, and these types of device are being used to treat hundreds of thousands of patients. Global drug-delivery systems sales were in excess of US$140 billion in 201166, and the global biomaterials industry is estimated to reach US$50 billion this year67, illustrating the potential economic impact of a new generation of devices. The annual number of medical implants — devices that replace some or all of a biological structure — is growing concurrent with our increased lifespan. For example, the combined number of hip and knee operations exceeded 1 million in the United States in 2010 and continues to grow68. Despite the huge clinical and commercial success of classical devices that provide solely a mechanical function, challenges with their longevity and ability to fully integrate in the host have persisted. Macroscale drug-delivery technologies have already had considerable clinical impact as they can aid favourable host integration and inhibit undesirable tissue responses through controlled delivery of select bioactive agents. Many more technologies are in preclinical and early clinical trials. We begin by describing the integration of MDD technologies with medical devices whose primary function is to provide structural support to local tissue (for instance cardiovascular stents), and then describe the role of MDD technology in orthopaedics. Drug-eluting stents. The epidemic of coronary artery diseases in the western hemisphere has led to widespread use of stents (some of which incorporate MDD technologies) to reperfuse coronary vessels. Although bare metal stents offer therapeutic benefit, ~25% fail through restenosis and thrombosis69,70. Drug-eluting stents (DES) that combine conventional bare metal stents with polymer coatings that control release of anti-restenotic drugs have helped to combat this problem71. The first two DES to receive approval in Europe and the United Sates were the sirolimus-eluting stent (2003; CYPHER) and the paclitaxel-eluting stent (2004; Taxus; Fig. 5a). These two drugs have primarily immunosuppressive and antiproliferative effects72. Clinical trials showed a decrease in failure rate (~8%) for both stents, which was attributed to reduced neointimal hyperplasia69,70. But the non-degradable polymers used in these first-generation devices can themselves cause thrombosis, and newer strategies are being developed that use biodegradable coatings or entirely biodegradable stents, non-polymer coatings with nanoscale features to control release, and genetically engineered cells. Clinical trials comparing DES that use a biodegradable polymer coating (such as biodegradable polylactide) on a metal stent to permanent-polymer DES demonstrated equivalent efficacy, and by 4 years the biodegradablepolymer-coated stents resulted in fewer side effects associated with late-stage thrombosis73,74. It is believed that the biodegradable coating provides a short-to-medium term advantage of controlled drug release, whereas the exposure of the bare metal following polymer degradation is believed to decrease the long-term chronic inflammation that would otherwise result from a permanent polymer coating, and to improve endothelialization73,75. A fully resorbable DES (Absorb) has shown promising results in phase I clinical trials, and enrolment is underway for a larger clinical trial76. Preclinical NATURE MATERIALS | VOL 12 | NOVEMBER 2013 | www.nature.com/naturematerials



NATURE MATERIALS DOI: 10.1038/NMAT3758 trials of a nanoporous alumina-coated stent showed positive effects in a rabbit model, but complications associated with coating fracture were found in a porcine trial; clinical trials are pending77,78. A different approach is to seed endothelial cells that are transfected with a VEGF gene on a bare metal stent; preclinical tests showed reduced neointimal hyperplasia on implantation79. Orthopaedic integration. A second surgical harvest procedure is necessary in certain conventional orthopaedic procedures, but incorporating MDD technologies into implant design may avoid tissue harvest, donor-site morbidity and other drawbacks of autografts. The INFUSE Bone Graft device (Medtronic; Fig. 5b), which has FDA- and EMA- (European Medicines Agency) approval for certain spinal-fusion procedures, combines a prosthetic steel cage (LT-Cage lumbar tapered fusion device) with a collagen carrier that retards the release of recombinant human BMP-2 (ref. 80). The INFUSE system was further approved by the FDA and EMA for use in open tibial fractures, combining it with an intramedullary nail81. It also has FDA approval as a standalone device for sinus augmentation procedures and localized alveolar-ridge augmentations82. Devices that deliver recombinant osteogenic protein-1 (OP-1 or BMP-7) in collagen I, OP-1 Putty (Olympus Biotech)83, and recombinant human-platelet-derived growth factor (rhPDGF) in beta tricalcium phosphate (β-TCP; GEM21S, Osteohealth), are also approved. Despite the commercial and clinical successes, these protein therapeutics (that is, INFUSE and OP-1 Putty) have come under scrutiny because of off-label use and/or improper marketing techniques, with the growth of ectopic bone in the spinal canal being reported in certain uses84. Because the carriers do not provide tight control over BMP release, high doses of the drug are implanted, and the devices initially present supraphysiological BMP levels. Improved control over release may allow more physiologically consistent doses, as demonstrated by using different material systems85–87. An alternative approach is to mimic how nature patterns tissue formation using spatially separated production of both promoters and inhibitors. Diffusion of each towards the source of the other leads to spatially precise regions in which tissue is formed88. This concept has been explored by simultaneously providing antibodies to the therapeutic protein in the regions outside of the treatment zone, leading to highly precise spatial control89. Future therapeutics will clearly benefit from improved control over factor presentation, and this is likely to result from the optimization of material systems and/or the combination of stimulatory and inhibitory signals. Macroscale drug-delivery technologies are also showing promise in solving the challenge of implant integration in skeletal tissue. Preclinical studies of hip implants with coatings of collagen90, and layer-by-layer polymeric coatings that release BMP (ref. 46), have demonstrated increased bone ingrowth into the porous implant exterior. This approach may be useful for difficult arthroplasty cases and revision surgeries. Similar approaches are being investigated in periodontics to augment the alveolar ridge. Combining rhPDGF with a deproteinized bone allograft containing denture fixings can reduce the required number of surgeries from three (tissue harvest, vertical ridge augmentation and denture placement) to a single procedure91,92.

MDD in tissue engineering and regenerative medicine

Tissue engineering and regeneration strategies that manipulate the body’s natural repair process to drive functional tissue formation are creating a new approach to treating patients. Sales of tissue-engineering products were recently reported as US$3.5 billion per year, with ~14,000 people employed in the industry93. Macroscale drugdelivery devices are central to many tissue-engineering strategies in which they deliver a cellular payload and/or morphogens or growth factors that direct cell behaviour. In tissue-engineering strategies,

the carrier material is often referred to as a scaffold94. These devices control cell behaviour by recruiting, housing and differentiating the cells, by delivering cells to the implant site and stimulating tissue formation, or by stimulating outward migration of cells from the implant (Fig. 6). In the following we first describe MDD materials for delivering cellular payloads without growth factors, and then discuss the role of multifunctional growth-factor delivering MDD devices in tissue-engineering strategies and regeneration by providing representative examples.

Growth-factor-free MDD with cellular payloads. Many cell therapies have been tested that do not use a biomaterial carrier. A common theme to these clinical trials is significant cell loss (>90%) within hours of delivery, and engraftment of a very small percentage of the cells at the tissue site of interest95–98. Many cell types need to be anchored in order to avoid undergoing anoikis (anchorage-dependent programmed cell death) during or immediately after delivery. Macroscale drug-delivery materials can provide anchoring ligands for cell survival and, more broadly, provide an appropriate microenvironment to enhance the function of transplanted cells99–101. Furthermore, the biomaterial can control cell release, improve cell localization, mechanically stabilize the environment and guide the structure of engineered tissue. Both synthetic and natural polymers have been used as scaffold materials, with cell adhesion facilitated through protein adsorption on their surface, or through ligands that are naturally occurring or added during processing. For example, polylactide (PLA), polyglycolide (PGA), or their copolymer (PLG), are synthetic polymers that are known to adsorb serum proteins (such as fibronectin or vitronectin) that have ligands favourable for cell binding102. A variety of cell types, including hepatocytes, fibroblasts and chondrocytes, have been seeded on these matrices, cultured, and implanted to engineer tissues such as skin, liver and cartilage103–105. These polymers can also be functionalized with specific adhesion molecules, such as the ligand Arg–Gly–Asp (RGD), a common adhesion ligand on many ECM proteins106,107. A variety of natural ECM proteins have also been used to fabricate the matrix; for example, products a

b

Poly(styreneb-isobutyleneb-styrene) containing paclitaxel

Collagen/ rhBMP-2 affinity retards release

Strut

Figure 5 | Examples of multifunctional MDD materials. a, Taxus (Boston Scientific) is a paclitaxel-eluting stent, which uses a polymer (poly(styreneb-isobutylene-b-styrene)) coating to control drug release by retarding drug diffusion. b, INFUSE bone graft and LT Cage (Medtronic). INFUSE consists of a rhBMP-2-soaked collagen sponge that stimulates bone formation in spinal reconstruction procedures. Affinity and diffusion retard the release of rhBMP-2 from the sponge. Figure adapted with permission from: a, © 2013 Boston Scientific Corporation; b, © Medtronic.



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NATURE MATERIALS DOI: 10.1038/NMAT3758 c

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Figure 6 | Macroscale drug-delivery technologies play a central role in many tissue-engineering devices, as they deliver factors that direct cell behaviour in various strategies. Typically, the goal is to regenerate various tissues, including muscle, blood vessels, bone and cartilage and return them to a healthy state, as shown in the histological images. a, An MDD device can deliver factors to recruit, maintain and differentiate cells from the host environment whereas its carrier material provides a housing in which cells can attach and grow. b, For accelerated healing, or in environments with insufficient host cells to recruit, the MDD device can be seeded with cells before implantation. The factors released by the device to stimulate tissue regeneration affect both the implanted cells and the cells in the local environment. c, In cases where cells migrate from the carrier for integration with host tissue, the cells can either be seeded before implantation or recruited from the host environment. The roles of the factors in the MDD device include programming the cells and/or directing their outward migration. Scale bars in images of muscle, vessel and bone, 50 μm; cartilage, 100 μm. Image of cartilage reproduced from ref. 165, © 2011 NPG.

based on bovine collagen, such as INFUSE and Apligraf, have made it to the clinic for wound healing. Cell-adhesion ligands are typically contained within these natural matrices. Other natural polymers derived from non-animal sources (such as alginate) do not contain ligands to which mammalian cell receptors can bind, and may not thermodynamically favour protein adsorption. These properties can make these materials suitable in certain applications where cell adhesion is not required (see immunoisolation below)108,109, and in situations where one desires to engineer a highly specific mode of cell adhesion with minimal background cell binding. For example, alginate can be modified with covalently coupled cell-adhesion peptides to mediate cell binding through specific integrin receptors109. This type of injectable matrix has been used to transplant a variety of cell types (including osteoblasts and chondrocytes) to engineer bone and mimics of growth plates110,111. These matrices can be designed to fully encapsulate and isolate a cell from the surrounding tissue, or to have an interconnected microporous structure that aids migration of transplanted and host cells in and out of the matrix. Microporous scaffolds can be formed by a variety of techniques, including the incorporation of sacrificial porogens during fabrication, freeze-drying, fibre bonding, threedimensional (3D) printing or laminate moulding112–118. Sacrificial porogens (such as salt or sugar) have been used in the fabrication of PLG scaffolds, for example112,113. Open-pore, macroporous PLG scaffolds can be formed with this technique by exploiting the high solubility of carbon dioxide gas in PLG, and this allows one to avoid the use of organic solvents in the fabrication and improve the stability of incorporated cells and proteins112,113. Another example is the freeze-drying of collagen–glycosaminoglycan (GAG) scaffolds. A collagen–GAG/H2O solution is frozen at a constant cooling rate, with the protein pushed to the edges of ice crystals that are subsequently sublimated, giving a microporous architecture whose pore size is a function of the final freeze temperature114. When these collagen–GAG scaffolds were seeded with keratinocytes and implanted in full-thickness wounds in a porcine model, skin tissue with many normal structural components was formed by 4 weeks119. Fibrebonding has been demonstrated for PGA among other materials. A non-woven mesh of PGA was mixed with poly(l-lactide) (PLLA) and heat-treated to bond the fibres at the interface. The PLLA that maintained the meshwork structure subsequently underwent 1012

solvent dissolution to leave a porous network116. In one application, these PGA fibres were moulded to form a bladder, subsequently coated with PLG for mechanical support, seeded with primary muscle and urothelial cells, and used to generate a bladder in vivo120. In lamination moulding, layers of porous polymer are cut to a defined shape, stacked and chemically bonded together115 (for example PLLA and PLG with chloroform). Polymer 3D printing can be similarly performed by layer-by-layer inkjet printing of a binder (such as chloroform) on powder polymer117 (such as PLLA). Cells have also been printed in 3D matrices118. Encapsulated cells are typically delivered in nanoporous biodegradable hydrogels (such as hyaluronic acid, fibrin, PEG or gelatin) that allow for diffusion of nutrients and waste products to and from the cells121. These injectable materials are advantageous as they can often be designed to gel in situ, which means that they can flow into and conform to a defect, potentially improving integration. These systems can undergo a sol–gel transition at physiological states (in fact, collagen gels at body temperature). Gels can also be ionically crosslinked (for example Ca2+ crosslinking of alginate) or covalently crosslinked before injection. However, care must be taken to ensure that the crosslinking products are not cytotoxic preand post-gelation. Cell-delivering MDDs can be designed to undergo breakdown or to resist degradation. In fact, MDD degradation with generation of ECM by implanted and infiltrating local cells is often desired in tissue engineering. Stable devices are desirable if the purpose of the MDD is to protect the transplanted cells from the host environment (as in immunoisolation). Hydrolytic breakdown (of PLG, for instance) is commonly used as a degradation mechanism in synthetic polymers. Other carriers, particularly those based on natural polymers (such as collagen or hyaluronic acid) can undergo enzymatic breakdown. Natural polymers that are resistant to enzymatic degradation can be chemically modified to be susceptible to hydrolysis. Alginate, for example, can be made susceptible to hydrolytic breakdown by partial oxidation to create acetal groups that are susceptible to hydrolysis122. Alternatively, degradation of the material may be undesirable. In immunoisolated therapies, xenograft tissue or cells are encapsulated so that they are not accessible to the host immune system and can survive to provide chronic signalling123. For example, porcine islets have often been delivered NATURE MATERIALS | VOL 12 | NOVEMBER 2013 | www.nature.com/naturematerials


NATURE MATERIALS DOI: 10.1038/NMAT3758 in non-degradable alginate (which serves to protect them from immune attack). This can induce normoglycaemia in non-human primates for up to 2.4 years124. A similar approach has been investigated for xenograft hepatocyte transplantation125. Although this concept has been around for more than 50 years, and some clinical trials have taken place, the challenges of maintaining cell viability, isolating the cells from the immune system over the long term, and finding a suitable cell source still remain123. Cell recruitment, housing and differentiation. For wound sites with sufficient repair cells in the local environment, the typical role of MDD devices is to attract the cells, provide a template for them to differentiate, and control the generation of functional replacement tissue. Multipotent stem cells or progenitor cells, which are known to play a central role in tissue development and repair, are often the target cell populations of these devices126. Signalling from the MDD device may be essential to selectively directing the fate of these cells, given their multipotency. In bone-tissue engineering, for example, mesenchymal stem cells and osteoprogenitors can be recruited to polymeric scaffolds (such as biodegradable polyurethane/PLG systems127 and RGD-modified alginate128) via the release of rhBMP-2, successfully healing non-union fractures. Poly(bis(pcarboxy)methane anhydride) (PCPM; ref. 129) and poly(propylene fumarate) (PPF; ref. 130) have been used to control the release of demineralized bovine bone extracts and the osteogenic thrombin peptide, TP508, respectively, to promote osteogenesis. Articular cartilage is a significant target for tissue-regeneration strategies because of its susceptibility to acute injury and degeneration, as well as its limited ability for self-healing. Adult chondrocytes are generally quiescent in situ and/or can exhibit stress-induced senescence131. Hence, cartilage tissue-engineering strategies often focus on the recruitment of mesenchymal stem cells (MSCs) from the bone marrow, periosteum or synovium. A pathway for release of marrow MSCs to a cartilage defect can be created by perforating the subchondral bone plate132. Although this microfracture technique is used clinically in isolation, improved results have been demonstrated by providing a scaffold that creates a framework for cell attachment, and subsequent proliferation and deposition of ECM proteins133. TGF-β is known to stimulate MSC recruitment, as well as chondrocyte differentiation and subsequent cartilage matrix deposition134. A strategy that combines collagen scaffolds with fibrin to control TGF-β release is now in a clinical trial135. A poly(εcaprolactone) and hydroxyapatite scaffold infused with TGF-β3adsorbed collagen hydrogel has also been used to regenerate hyaline cartilage on the humerus in a rabbit model136. Although the cell source was not fully evaluated, it was proposed that stem or progenitor cells of synovium, bone marrow, adipose or even vasculature were recruited and subsequently differentiated to chondrocytes. An alternative strategy for cartilage-tissue engineering is to generate cartilage at an auxiliary site where MSCs or chondroprogenitors are available. For example, recruitment and differentiation of cells from the periosteal cambium layer has been demonstrated using a combination of polycaprolactone and fibrin that controls release of TGF-β in a subperiosteal pocket137. Trauma, burns or disease of the skin can lead to severe medical complications, creating a strong impetus for the development of new therapies. In diabetic patients alone, foot ulcer treatments and amputations were estimated to cost the US economy around US$38 billion in 2007138. One approved therapy for diabetic foot ulcers, Regranex, uses a methylcellulose gel to deliver rhPDGF to stimulate the recruitment of fibroblasts and smooth muscle cells, as well as ECM production139. But the gel provides little control over the release kinetics of the PDGF, and thus several systems that aim to provide better control over the release of PDGF or other factors

INSIGHT | REVIEW ARTICLES that affect wound healing (for instance TGF-β, FGF, VEGF and epidermal growth factor (EGF)) have been tested140. INTEGRA, an FDA-approved bilayer collagen–GAG/silicon skin-regeneration template141, has been combined with different factors that actively stimulate cell recruitment and matrix formation, and tested in preclinical wound-healing models. TGF-β1 delivery accelerated wound contraction and epithelialization142. Accelerated healing was also found when INTEGRA was doped with bFGF-containing gelatin microspheres143. Preclinical studies in which an immobilized adenoviral vector for PDGF was delivered from a collagen gel also demonstrated efficacy144. Cell-containing implants for local tissue regeneration or hormone replacement. Macroscale drug-delivery devices may be used to promote or accelerate healing of large wounds by delivering and maintaining cells that have been genetically modified to overexpress and secrete select factors, or by delivering a combination of cells and bioagents. The cells can be autologously sourced, with or without culture expansion, or can be allogeneic. PLG scaffolds delivering condensed plasmid DNA encoding for BMP-4 (to provide an osteogenic stimulus) and VEGF protein (to enhance angiogenesis) improved bone formation from transplanted human bone marrow stromal cells (hBMSCs)145. MSCs from marrow, muscle and adipose tissue were transfected ex vivo to overexpress BMP, and implanted in scaffolds to demonstrate their osteogenic potential146–148. Ex vivo transfection of stem cells from bone149, adipose tissue150 and muscle151 to overexpress VEGF before their implantation in a scaffold enhanced angiogenesis and bone formation in preclinical models. Similar approaches to scaffold vascularization have been shown in preclinical models of other tissues. For example, VEGF-transfected myoblasts were delivered in polyurethane for cardiac repair after myocardial infarction152, and a MSC line that naturally expresses high levels of VEGF was delivered in collagen scaffolds to repair full-thickness murine skin wounds153. Non-transfected cells, delivered in combination with a protein drug, have shown success in an ectopic implant model for bone (alginate gels delivering hBMSCs along with rhBMP-2 and TGF-β3; ref. 85) and in cartilage models (hydrogels of oligo(poly(ethylene glycol) fumarate) containing TGF-β1-loaded gelatin microparticles and bone MSCs154). Severe skeletal muscle injuries can be debilitating. Surgisis — a cellular pig jejunum containing angiogenic and other factors — was used to culture endothelial cells, myoblasts and foreskin fibroblasts for 3 weeks154. These scaffolds were subsequently implanted in a skeletal muscle wound, and demonstrated vascular integration with the host and functional muscle formation within the scaffold by 2 weeks155. Tissue-engineering strategies156,157 and immunoisolated xenografts123–125 have been explored to replace hormone secretion from defective or missing glands. Genetic modification of different tissue structures (that is, non-glandular) to replace lost hormonal secretion has also been proposed158. For example, post-mitotic myofibres were induced to form a muscle-like structure in vitro and genetically modified to overexpress rh-growth hormone (rhGH). Following subcutaneous implantation, these maintained pharmacological rhGH levels in sera for up to 12 weeks158. Stimulating outward migration of cells from the implant. Treatment with MDD devices that release cells into an injury site may be able to reverse the typical course of inflammation and cellular apoptosis, and effectively promote regeneration in certain types of tissue injury. In skeletal muscle, for example, macroporous alginate scaffolds delivering VEGF, IGF-1 and myoblasts to a severely wounded skeletal muscle led to a marked increase in myoblast engraftment, limited fibrosis and increased function of regenerated muscle tissue, as contrasted to cell delivery alone159. Similarly,



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REVIEW ARTICLES | INSIGHT a macroporous alginate carrier was used to deliver VEGF along with endothelial progenitor cells into a murine hindlimb ischaemia model that typically results in autoamputation in just 3 days160. Mice treated with the combination of cells and VEGF regained full limb perfusion by day 40, and had greatly diminished tissue loss160. Whereas tissue regeneration has been the focus of MDD devices so far, MDD may also be highly relevant to immunotherapies. Cellbased cancer-vaccination approaches have been developed, but these typically necessitate cumbersome and expensive ex vivo cell manipulation and transplantation. Recently, devices based on MDD technologies have been described that instead perform the immune cell manipulations in situ161. A device based on PLG was designed to simulate an infection by releasing granulocyte–macrophage colony-stimulating factor (GM-CSF) to recruit and proliferate host dendritic cells, by releasing tumour lysates (which are processed by dendritic cells and presented to other immune cells), and by presenting cytosine–guanosine oligodeoxynucleotide (CpG-ODN) sequences to mimic bacterial DNA. These devices recruited large numbers of antigen presenting cells, promoted their activation and led to their efficient trafficking to the lymph node to stimulate an antitumour T-cell response161. This approach led to complete regression of established melanoma in preclinical models162. The concept of recruiting and programming cells to affect biology at a remote site has potential in a number of applications.

Outlook

There have been remarkable advances in the field of MDD over the past 45 years. Millions of patients benefit annually from the therapeutics that have arisen from this field. The improved spatiotemporal control of drug presentation that is possible using these drug-delivery devices has increased drug effectiveness and reduced side effects, and enabled the delivery of sensitive protein therapeutics for extended time periods. Macroscale drug-delivery as a component of multifunctional devices is leading to new therapeutic approaches to disease and injury in the medical-device and tissue-engineering fields. Taking vascular stents as an example, the incorporation of a MDD coating or component was rapidly taken up by the clinical community, in part because the core principle of the device was unaltered. Many more surgical implants may benefit from MDD strategies that assist integration, control inflammation or prevent infection. Despite the existence of over 25,000 papers describing successful implementation of MDD technologies (according to a search for ‘Controlled drug delivery’ in PubMed), challenges in the clinical translation of these devices persist. One hurdle is the acceptance of multifactor devices by industry and the various regulatory bodies. These devices have complex functions and non-traditional fabrication processes. Because of rapid progress in this field, there is often neither consensus on the best preclinical tests to perform nor on how to classify and approve these devices. Another challenge is the high cost associated with the active agents, which are often recombinant proteins and cells. Because of all this, the bench-to-bedside translation period is often extensive and expensive. Multiple aspects of the clinical translation during early product development have been addressed, including strategies for in vivo delivery, sterilization, minimizing in vitro manipulation and preparing devices that augment repair beyond drug delivery alone (or at least that diminish side effects). But further work is necessary to develop a larger catalogue of materials that better address these issues as well as other challenges that have arisen over time. For example, in tissue engineering and regenerative medicine, developing ‘off-the-shelf ’ devices that minimize or eliminate the use of expensive and prolonged in vitro culture and intraoperative preparation is likely to improve clinical acceptance. Fabrication methods are still needed that are harmless to the bioagent and/or cells while 1014

NATURE MATERIALS DOI: 10.1038/NMAT3758 minimizing their loss prior to delivery and that are amenable to terminal sterilization. There has also been a push to minimize the invasiveness of delivery, with injectable MDDs being favoured when invasive surgery is not otherwise needed. Although success has been demonstrated with several injectable materials, many of these have significant drawbacks (such as poor mechanical properties and potentially harmful byproducts of crosslinking). New crosslinking systems and fabrication methods are likely to help broaden the catalogue of MDD materials. Finally, the design of devices that allow on-demand release of multiple bioagents is still in the developmental phase, and designing MDD devices to favourably affect the biology of the target cells and tissues, while providing the desired drug release, may be critically important in a number of applications (for example in the delivery of chemotherapy agents that modify cancercell responsiveness in the context of stromal signalling). We expect that ever-increasing knowledge of disease combined with data on MDD devices can provide therapeutic effectiveness and create incentives to overcome current challenges. The next decade will also continue to see an explosion in our understanding of the biology underlying disease, and in the identification of new therapeutic targets that will often require MDD technology. There will be continued exploration of the science behind the current generation of MDD therapeutics, optimization of their delivery patterns and further integration of MDD into medical devices. The exploration of small molecules (such as peptides and oligonucleotides) capable of replacing functions intrinsic to larger molecules may reduce some of the challenges faced in the development of MDD devices. Innovative institutional structures that bridge academia and industry to aid the clinical translation of devices may also play a key role. As a result of this confluence of activities, the safe clinical implementation of effective MDD devices is expected to accelerate. Received 18 February 2013; accepted 15 August 2013; published online 23 October 2013

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Acknowledgements

Funding from the NIH to the authors’ laboratory is gratefully acknowledged.

Additional information

Reprints and permissions information is available online at www.nature.com/reprints. Correspondence and requests for materials should be addressed to D.J.M.

Competing financial interests

The authors declare no competing financial interests.



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