Sustained release emphasizing recombinant human bone morphogenetic protein-2

Advanced Drug Delivery Reviews 31 (1998) 303–318

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Sustained release emphasizing recombinant human bone morphogenetic protein-2 a b a, Shelley R. Winn , Hasan Uludag , Jeffrey O. Hollinger * a

Division of Plastic and Reconstructive Surgery, Oregon Health Sciences University, 3181 SW Sam Jackson Park Road, Portland, OR 97201 -3098, USA b Genetics Institute, One Burtt Rd, Andover, MA 01810, USA

Abstract Bone homeostasis is a dynamic process involving a myriad of cells and substrates modulated by regulatory signals such as hormones, growth and differentiating factors. When this environment is damaged, the regenerative sequalae follows a programmed pattern, and the capacity for successful recovery is often dependent on the extent of the injury. Many bony deficits that are excessively traumatic will not result in complete recovery and require therapeutic intervention(s) such as autografting or grafting from banked bone. However, for numerous reasons, an unacceptably high rate of failure is associated with these conventional therapies. Thus, alternative approaches are under investigation. A class of osteogenic regulatory molecules, the bone morphogenetic proteins (BMPs), have been isolated, cloned and characterized as potent supplements to augment bone regeneration. Optimizing a therapeutic application for BMPs may be dependent upon localized sustained release which in kind relies on a safe and well characterized carrier system. This review will discuss the current status of BMPs in bone regeneration and specifically will present the potential for a clinical therapeutic role of recombinant human BMP-2 sustained release carrier systems.  1998 Elsevier Science B.V. Keywords: Bone regeneration; Collagen; Growth factors; Osteoconduction; Osteoinduction; Poly(a-hydroxy acids); rhBMP2; TGF-b superfamily; Tissue engineering

Contents 1. Introduction ............................................................................................................................................................................ 2. Bone morphogenetic proteins (BMPs)....................................................................................................................................... 3. Carrier strategies ..................................................................................................................................................................... 3.1. Poly (a-hydroxy acids) (PHAs) ......................................................................................................................................... 3.2. Collagen.......................................................................................................................................................................... 3.3. Alternative carriers........................................................................................................................................................... 4. Tissue engineering strategies .................................................................................................................................................... 5. Clinical applications ................................................................................................................................................................ 6. Conclusions ............................................................................................................................................................................ References ..................................................................................................................................................................................

*Corresponding author.

0169-409X / 98 / $19.00  1998 Elsevier Science B.V. All rights reserved. PII S0169-409X( 97 )00126-9

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1. Introduction

2. Bone morphogenetic proteins (BMPs)

The capacity for damaged bone to repair itself, provided the damage does not result in a defect of a ‘critical size’, has been an integral component of vertebrate evolution. In cases of trauma resulting in critical-sized bony deficits, however, the damaged bone will not regenerate spontaneously, thereby requiring either autografts or allogeneic banked bone to restore form and function (reviewed in Refs. [1,2]). These procedures are performed in the range of 200 000 times annually in the United States [3], with an unacceptably high rate of failure for autografts (13–30%, [4]) and banked bone (20–35%, [5]). Unacceptable clinical outcome, as well as issues associated with potential pathogenic transmissions and immune responses in allografts [6–9], warrant the development of safe and efficient alternatives. In order to arrive at a clinically relevant alternative to conventional bone repair therapies, it is critical to understand the bone repair process, which can provide a framework for mimetics. Systematic analyses of fracture repair [10–14] have established the sequalae of events and interactions between the critical physical, molecular, and cellular elements responsible for repair. Engineering scaffolds for bone regeneration requires a mimetic that will provide an environment for cell recruitment, attachment, expansion, differentiation, and the interactions of the cellular elements with extracellular matrix proteins [15,16]. A structure or scaffold for bone regeneration is pivotal [17]. In addition, endogenous molecules may be equally important to ensure bone regeneration. Indeed, it had been speculated that ‘‘some specific bone formation substance activating the nonspecific mesenchymal tissue’’ was a critical element necessary for bone regeneration [18]. The pursuit to identify this specific bone formation substance initially led Professor Marshall Urist to coin the term autoinduction [19], which was later adapted to osteoinduction, defined as the ability to induce bone formation at an ectopic (non-bone) site [20]. The osteoinductive class of molecules, specific to bone formation substances critical for bone regeneration, were termed bone morphogenetic proteins (BMPs) [21].

Determining functional roles of BMPs has sparked a plethora of investigations into their relationship in bone physiology and their potential as therapeutics for bone regeneration. BMPs are categorized within the transforming growth factor (TGF)-beta (b) ‘superfamily’ [22–25]. This superfamily is represented by numerous molecules, and has recently been reviewed [22,23], but is nevertheless an evolving class of molecules whose list continues to expand [26,27]. The BMPs within the TGF-b classification include BMPs 2–15 (BMP-1 is not part of the TGF-b superfamily) [27,28] (information for BMP14, J. Wozney, personal communication; BMP-15 [29]). Deciphering the sequences of BMPs 2–9 for identification was primarily the result of efforts of Wozney et al. [26–28], Wang et al. [30] and Celeste et al. [31] utilizing oligonucleotide probes to screen genomic libraries. Recently, the BMPs 10–15 were identified with hybridization and polymerase chain reaction technology [32–34]. A thorough presentation of the BMPs, in terms of their properties, roles, locations and terminology, can be reviewed elsewhere [35]. The focus of this article will be centered on the delivery of BMP-2, more specifically, recombinant human BMP-2 (rhBMP-2). BMPS are prominent molecules necessary for critical path development of tissues and organs early in embryogenesis, that continue post-fetally, and are necessary for stimulating bone growth and morphologic differentiation in species as disparate as Xenopus to Homo sapiens [36]. Later in life, BMPs are sequestered primarily in bone and are liberated following a fracture or other bone damage, recruiting pluripotential cells to take residence and differentiate into osteogenic cells [37–39]. BMPs are not localized as discrete entities, rather they exist as a milieu interacting as vital components for bone maintenance and repair [40,41]. Additionally, BMPs have been utilized experimentally to induce the differentiation of pluripotential cells residing in muscle toward cartilage and bone nodules [37,42]. Thus, BMPs have been shown to be potent soluble factors impacting the musculoskeletal system, primarily targeting bone regeneration in animal studies [43], and recent-

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ly showing great clinical potential [44]. With this potential in hand, the ultimate goal for BMPs would be to develop human therapies that may someday replace autograft and allogeneic procedures. It is emphasized, however, that the safe and practical delivery of BMPs may be realized in a sustained delivery system. Ideally, the delivery system should provide several functions: a physical scaffold to provide a substrate for replacement bone cells and, secondly, the sustained ability to actively deploy BMPs for cell interactions. The present review will focus on two polymer systems, those based on a class of the poly(a-hydroxy acids), and secondly on collagen(s).

3. Carrier strategies A wide range of biomaterials have been utilized as osteoconductive scaffolds, i.e. scaffolds which allow regeneration and ingrowth of bone tissue at osseous sites. Some of these scaffold biomaterials include bovine collagen [45,46], demineralized bone matrix [47–49], calcium phosphate ceramics [50,51], bioglasses [52], organoapatites [53,54], polylactide homopolymers [55–57], poly(lactide-co-glycolide) polymers [58–60], polyanhydrides [61], organophosphates [62–64], and polyphosphazenes [65]. These scaffolds are designed to exploit the inherent regenerative capability of osseous tissue by allowing bone marrow, periosteum or mesenchyme-derived osteoprogenitor cells to populate and mineralize the scaffold. Osteoconductive biomaterials can serve as a starting point for the design of BMP carriers. Further optimization of carrier properties should be considered since the carrier properties critical for osteoconduction may not be optimal or even compatible for osteoinduction (bone formation at an ectopic site). Obviously, each class of osteoconductive biomaterials has its advantages and disadvantages and other reviews have provided detailed accounts of osteoconductive biomaterials used as BMP carriers [56,66]. It is tempting to compare the potency of different osteoconductive scaffolds as BMP carriers but the variations in preclinical models (differences in animal species, age and anatomical sites of implantation, etc.) and the BMPs (especially for purified BMP

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preparations) make this a difficult task. However, a limited number of reports, in which the performance of various BMP carriers were tested head-to-head in the same animal model [54,67–69], may provide significant observations to lay a foundation for the design of optimal BMP carriers. Optimizing the clinical application for BMP delivery may be dependent upon localized sustained release. Preclinical studies have stressed this issue based on the concepts that sustained release at low levels may be more efficacious than a bolus load and additionally may pose less risk from unanticipated side-effects [70,71]. Foremost, a clinically-relevant carrier system must be safe and, secondly, must exhibit properties which are well characterized and, therefore, reproducible. In addition, anatomical differences will require the ability to modify the form of the carrier, e.g. differences in blood flow will impact the concentration gradient of released rhBMP-2 [35]. Several candidate materials may fulfil the criteria necessary for a successful carrier: the first being the class of polymers known as the poly(ahydroxy acids) (PHAs), and another represented by collagens, a logical choice, since the major protein component of bone is type I collagen [38]. Each of these candidate materials will be discussed for their potential to deliver rhBMP-2 for bone repair. Based on the extensive experience generated in preclinical animal studies, collagen (and DBM) has provided a satisfactory delivery system for rhBMP [67,68,73–77]. Moreover, collagen has repeatedly been observed to be an effective carrier for rhBMP, independent of the site of implantation. Collagen offers cells a permissive substratum for attachment and differentiation [43,72] which may be a result of collagen’s composition as the major component of the extracellular matrix (ECM). The ECM influences the biology of cells through cell-surface receptors that interact with growth promoting factors, such as BMPs, thus providing an interactive environment [16,43,72,78,79]. Both the biologic requirements, as well as the engineering parameters, require careful consideration when defining optimal characteristics for carrier systems. Some of the critical characteristics to consider will include surface chemistry, surface charge, surface texture, pore size range, void volume, biodegradation rates, hydrophobicity and

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hydrophilicity, crystallinity, and release kinetics of incorporated molecules such as growth factors [56,57,80,81]. Another issue that remains to be discussed is sustained release. Is sustained release preferred over release of a pulsatile or bulk nature? Is BMP required throughout the bone regeneration cascade and what are the optimal doses? These issues, of course, remain to be elucidated. The type of rhBMP2 release dynamics necessary for enhanced bone healing may not be completely satisfied by PHAs or collagen. Presently, both of these materials are insufficient to allow coupling with growth or differentiating factors, such as BMPs, that may result in sustained release for durations of several weeks or months. If such pharmacokinetics of BMP release are desired, innovative technologies will need to be incorporated with scaffold engineering, resulting in sustained release that otherwise would not be possible.

3.1. Poly (a -hydroxy acids) ( PHAs) The specific PHA homopolymers of polylactide (PL), polyglycolide (PG), and their copolymers of poly(lactide-co-glycolide) (PLG) have had extensive biomedical applications. As suture materials, PG and PLG have over a 25 year history as clinically safe and efficient biomaterials. In addition, these materials are biodegradable (reviewed in Refs. [56,57,80]). Biodegradation is initiated by contacting biologic fluids in situ, where water elicits non-specific hydrolytic cleavage, which ultimately results in the formation of lactic and / or glycolic acid. Both lactic and glycolic acid are natural molecules that can be metabolized by endogenous systems [82,83]. Many factors influence the rate of biodegradation and also impact the safety of the material [57,58,82]. Some of these factors are inherent to the materials (bulk properties), while others may evolve from the processes involved in their fabrication. For example, inherent characteristics of polymers determine the wettability of the material [82]. Since PG is more hydrophilic (water attracting) than PL, its biodegradation is more rapid due to the increased access of biologic fluids. This characteristic may also influence the attachment of cells. In general, the electronegative nature of cells prefers an electropositive and

‘wettable’ substratum for interactions and attachment [17]. However, as is common in biologic systems, there are exceptions to this generalization: many cells will not interact with untreated polystyrene (electropositive). Other characteristics of the material(s) may be unwanted contaminants, residuals or byproducts from the polymerization process [56]. Many residual molecules from the polymerization process, such as catalysts and solvents, as well as variations in crystallinities within the same polymer, can influence both the biocompatibility and biodegradation. Several publications are referenced that thoroughly review these issues in detail [56–58]. The biodegradation of these suture materials has also been reported to result in an outcome that is biocompatible (i.e., clinically safe) to the host; that is, the ability of the material to perform with an appropriate host response that promotes the desired effect without inducing harm to the host [84]. However, PHA suture biocompatibility does not ensure that PHAs formulated as bulk devices, such as carrier systems for BMPs, will elicit a similar response. Several items in particular have received attention concerning the biodegradation of devices for orthopedic applications. These items include the proposed in vitro decrease in local pH [85], the shedding of fragments during degradation [86] and adverse host tissue reactions that may result from a number of contributing factors [86,87]. Obviously, the presence of abundant macrophages and foreign body giant cells can negatively influence tissue repair / regeneration and should be avoided. For a biodegradable carrier to be successful, the carrier should not only be biocompatible, but should support bone ingrowth and the degradation rate should coincide with the rate of bone formation. Nevertheless, carrier systems designed from a material that has regulatory promise is advantageous for this class of the PHAs. Polylactide (PL) and polyglycolide (PG) synthesis is typically a result of ring-opening of the appropriate six-membered lactone monomer (i.e., lactide and glycolide) [88]. Lactide and glycolide can be prepared by dimerization of hydroxy acid precursors, lactic and glycolic acids, respectively [88]. Chiral chemistry impacts the carbon of lactic acid (LA), therefore it may polarize light to the left (levorotatory, or l) or right (dexterorotatory, or r). Glycolic

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acid (GA) is optically inactive since it lacks a chiral center. The configurations of optically active molecules are standardized against glyceraldehyde, which has two possible structures (enantiomers) for LA, identified by upper case L and D [89]. Many combinations are possible, which have been previously reviewed [56,57], as well as descriptions of the types of synthetic processes available to date [88]. Some of our recent efforts have utilized the racemic lactide of D,D-L,L-polylactic acid optimized to simulate the architecture of cancellous bone, referred to as OPLA , an acronym for Open-cell PolyLactic Acid [90] (Fig. 1). We have emphasized homopolymers for designing carrier systems, which undoubtedly result in a more homogeneous and, therefore, predictable outcome relative to the various combinations of lactide and glycolide precursors. There is an abundance of literature making observations with inadequately characterized homopolymers or copolymers that likely contribute to... ‘‘discrepancies between data obtained apparently under similar conditions’’... in wound healing studies [58]. In addition, controversy related to the safety of PLA and PGA as bone implants has been discussed, emphasizing the impact of the structure and molecular weight of the materials [91], as well as the use of different model systems [58,91]. One of the authors (Hollinger) has previously made some suggestions concerning consistencies in polymer nomenclature and techniques to adequately define the PHAs chemical structure(s) to avoid complications associated with the use of inadequate materials [56]. Other characteristics of the carrier dimensions will have an impact on the rate of polymer degradation and influence the pharmacokinetics for drug delivery. Determining and predicting the biodegradation rate of a PHA suture is straightforward relative to the complex structure involved in a carrier mimetic. Independent of structure, the rate of PHA biodegradation is influenced by the inherent property of bulk erosion [92]. Can polymers that are bulk eroders (mass loss) provide a stable scaffold for host attachment and differentiation to replenish endogenous bone that coincides with the rate of degradation? Additionally, to mimic the structure of bone, the internal architecture must be porous, which is defined in terms of pore size, pore distribution, void volume and voids that are interconnected [56]. The

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Fig. 1. Scanning electron micrographs of: (top) Open-cell PolyLactic Acid (OPLA ) macrostructure that has been manufactured to simulate cancellous bone. OPLA is a racemic PLA that is specifically D,D-L,L-Polylactic Acid (see text for details). The bottom micrograph represents a candidate collagen delivery scaffold for BMP, which is highly fibrous, and exhibits a smallersized void range than OPLA . Scale bar 5 100 mm. OPLA . Kindly provided by Dr. John H. Brekke, THM Biomedical, Inc., Duluth, MN.

importance of adequately reproducing the bone architecture cannot be overemphasized. Recent studies have observed that monolithic disks of PHAs combined with either partially-purified allogeneic or xenogeneic BMP were ineffective to regenerate

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osseous defects [93,94]. In contrast, rhBMP-2 and allogeneic bone collagen together have regenerated form and function in model systems of dogs [73] and rats [74,75]. It has been suggested that monolithic devices with inappropriate internal architecture may impede osteoconduction [56,81,95]. The interactions of a PHA carrier with host elements (cells and fluids) can be manipulated to modify the rate of biodegradation by changing the surface area of the voids. Materials such as calcium phosphates and polyethylene materials have been evaluated for their osteoconductive (bone ingrowth) capabilities. The nominal range of porosity for bone ingrowth was observed at 200–400 mm [51,96–99]. Similarly, this range was suggested for PHAs [81] and a recent study by Robinson et al. has evaluated the porosity of a PHA for osteoconduction [100]. Calvarial wounds implanted with racemic D,L-PL devices with 300–350 mm sized pores and a 60% void volume were observed to promote osteoconduction in contrast to smaller sized pores and void volumes that were unsuccessful [100]. An appropriate pore size range for osteoconductive carriers may be broadly acceptable (200–400 mm) but may require additional refinement with advancing technologies. Kenley et al. [101] have recently evaluated porous particles of PLG (1:1 molar ratio of lactide:glycolide), with a void volume of | 60%, and having a particle size range of 75–250 mm. Utilizing the critical-sized rat calvarium model, PLG (and a control insoluble collagenous bone matrix (ICBM)) was mixed with a known volume of biopolymeric ‘thickening agents’ [101]. For this study, the thickening agents allogeneic blood, hydroxypropyl methylcellulose or calcium crosslinked sodium alginate were loaded with either 0, 10 or 30 mg rhBMP-2 and inserted into critical-sized intraosseous defects. The 30 mg rhBMP-2 dose was shown to be equivalent to the ICBM group, generating additional data that refuted a report indicating that a collagenous substrate was required for effective osseous regeneration with the delivery of osteoinductive proteins [102]. However, the consistency of the materials resembled soft tissue and mechanical manipulations at the wound bed would often result in displacement. Lee and colleagues evaluated the healing of large segmental defects in rat femurs with non-porous small and large PLG particles loaded with low (0.93 mg),

medium (3.1 mg) and high (9.3 mg) doses of rhBMP2 [60]. The PLG particles were mixed with the various doses of rhBMP-2 in allogeneic blood and allowed to clot. The smaller particle size with the highest rhBMP-2 dose resulted in regenerated bone achieving the greatest strength (torsional stiffness and strength). However, as was observed in the Kenley et al. study, the consistency of the mixtures was typically inadequate. Other studies have utilized the particulate PLG-autogenous blood delivery system to deliver rhBMP-2 in a rabbit critical-sized radius model [103,104] and maxillary alveolar cleft repair in dogs [104,105]. The observations made from these studies also exhibited great promise for these types of therapies; however, an emphasis on improving the delivery of rhBMP-2 was stressed. Clearly, improvements in carrier vehicles are warranted. The development of ‘off-the-shelf’ devices for clinical application in osseous repair continues to be pursued. Sandhu et al. have evaluated rhBMP-2 delivered with an open-cell D,D-L,L-polylactic acid (OPLA ) biodegradable scaffold in a canine posterolateral (transverse process) spinal fusion model [106]. Each OPLA strip (12 3 6 3 30 mm) was soaked with the cargo capacity, 1 cc, of a solution containing 2.3 mg of rhBMP-2. Observations at 3 months indicated that autografts had resulted in no union, while the rhBMP-2 group produced 100% solid transverse process fusion(s). As a follow-up to the feasibility study, in which a massive 2.3 mg (2300 mg) dose was used, a dose–response study was undertaken to evaluate 58, 115, 230, 460 and 920 mg [107]. As was observed with the previous 2300 mg rhBMP-2 study, all rhBMP-2 study animals at 3 months exhibited intertransverse process fusion, in contrast to the autograft or carrier alone groups, which exhibited no fusion. Additional studies will be useful to determine the minimum threshold dose for inducing fusions. These studies have shown that a more manageable device, both in terms of handling and ‘off-the-shelf’, represents a great potential for biodegradable PHA carriers to be further developed, refined and tested in pre-clinical models.

3.2. Collagen Collagens are natural candidates for the delivery of BMPs since the major protein component of bone

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is type I collagen [38]. The collagens are typically derived from bone (whose mineral component is removed), skin or tendon. Demineralized bone matrix (DBM) was the initial carrier of choice for implantation of purified or recombinant BMPs. DBM is comprised of mostly insoluble, highly crosslinked type I collagen, although other non-characterized collagens and non-collagenous proteoglycans may be present. Guanidine / urea extracted DBM (from which soluble bone morphogenetic activity is removed) can serve as an excellent osteoconductive template by itself in bridging small bony defects. Since DBM is a bone-derived scaffold, it can provide an ideal template for the attachment, and differentiation of osteoprogenitor cells. When implanted with BMP, DBM can potentially affect the osteoinductive activity, thus participating directly in the mineralization process. For example, osteopontin and related phosphorylated sialoproteins may act as nucleators and / or inhibitors of calcium phosphate formation [108]. The presence of such proteins in DBM preparations may exert a synergistic and / or antagonistic effect to BMPs pharmacological activity. DBM used in preclinical studies are typically derived from an isogeneic source to eliminate immunogenicity issues of the implant. This was based on previous studies by Sampath and Reddi [109], who observed detrimental effects of allogeneic carrier immunogenicity on BMP induced bone formation. In a clinical setting, however, allogeneic DBM may elicit an inflammatory or immunogeneic response. Other potential concerns with DBM include the feasibility of disease transmission with allografts [6–9] and, as a particulate carrier, it is critical for DBM particles to be retained at the implant site to prevent bone formation at undesirable sites. Limitations of DBM have inspired the exploration of reconstituted collagens from other sources as alternative BMP carriers. Different types of collagens have been purified on a large scale [45,46]. The purification process typically yields a collagen solution or dispersion which can be fabricated into macroporous scaffolds in an implantable format (Fig. 1) [110]. Macroporosity is important for cellular access, since BMP action is based on the cell invasion of a carrier, hence the need for a freely accessible space within a carrier. The manufacturing process also may allow control over the wettability, mechanical properties and geometry (pore size and

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connectivity) of the carriers. Reduction of collagen immunogenicity has been observed following enzymatic treatment, e.g. by pepsin, to remove the collagens telopeptide fraction. In addition, both physical (e.g., dehydrothermal, UV-irradiation) and chemical crosslinking strategies (e.g., aldehydes, carbodiimides, etc.) have been explored to minimize immunogenicity and to provide control over in vivo biodegradation [111,112]. It is noteworthy that the in vivo cellular reaction and biodegradation of collagen scaffolds are different when they are implanted with BMP as opposed to implanted alone. Collagen sponges currently are being evaluated in both preclinical and preliminary clinical studies for rhBMP-2 delivery. Preliminary clinical results of rhBMP-2 delivered with absorbable collagen sponges for maxillary sinus floor augmentation have been presented [44]. The observations provided evidence that this combination of rhBMP-2 with absorbable collagen sponge was efficacious and therefore may be an acceptable alternative to traditional treatments for maxillary sinus floor augmentations. Successful and reproducible bony bridging is demonstrated in numerous critical size defect models, including goat maxillary sinus floor augmentation model [76], as well as canine periodontal [68] and mandibular [113] defect models. Implanting rhBMP-2 with collagenous carriers in numerous anatomical sites has clearly demonstrated that the pharmacological effects of BMPs were local. Once released from the implant site, the BMP, similar to TGF-b [114], is rapidly cleared from the circulation by the liver (unpublished observations), resulting in no activity at sites other than the implant site. To better understand the factors contributing to local osteoinduction, and specifically to evaluate the impact of rhBMP-2, the local pharmacokinetics of rhBMP-2 released from collagenous carriers were characterized. A well-established implantation model, the rat ectopic assay, was employed for this purpose [115]. The rhBMP-2 was the main protein used in this work but other natural and chemically modified rhBMPs were also used. Radiolabeled tracer techniques (both 125 I- and 35 S-methionine labeled) had to be used to quantitate in vivo levels of the proteins retained within the carriers since quantitation of cold BMPs is not feasible due to assay interference. The carriers were typically soaked with radiolabeled protein solutions (hot:cold ratio | 1:1000) and implanted as

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previously described [68,76,113]. After sacrifice, the implants were recovered, the radioactivity was determined, and was used as a measure of rhBMP-2 retained in the implants. The pharmacokinetics of rhBMPs were generally in two phases: an early rapid rhBMP-2 loss (within a few hours of implantation), followed by a more gradual rhBMP-2 loss. Typical rhBMP-2 pharmacokinetics from a reconstituted, chemically-crosslinked collagen carrier and a DBM particulate carrier are shown in Fig. 2A. A significant difference between the two carriers was observed initially, DBM retaining less than half of the dose retained by the collagen sponge. The subsequent rate of rhBMP-2 loss was monoexponential, irrespective of the nature of the carrier, and the concentration of the implanted rhBMP-2 solution (between 0.08 and 2 mg ml 21 ) did not affect the overall protein pharmacokinetics [115]. The sponge properties, however, were found to contribute to rhBMP-2 pharmacokinetics. In one series of experiments, where the collagen sponges were treated with formaldehyde (for crosslinking) and ethylene oxide (for sterilization), untreated sponges had a higher fraction of initial retention, presumably due to higher affinity for rhBMP-2, but exhibited a faster rate of rhBMP-2 loss as compared to chemically treated sponges (Fig. 2B). The relatively rapid rate of rhBMP-2 loss in the untreated sponges may result from rapid degradation. In addition to carrier properties, the intrinsic properties of the protein (such as pI) also affected the protein pharmacokinetics. For example, succinylated rhBMP-2 (rhBMP-2 whose free amino groups were converted into carboxylic groups by succinic anhydride; a pI shift from | 9 to | 3) exhibited a rapid loss from the implant site as compared to native protein (Fig. 2C). The rapid loss is attributed to lower affinity of the carrier for the succinylated protein and / or higher solubility of the succinylated protein in the biological milieu. These preliminary studies have only recently begun to provide clues about factors contributing to the in vivo pharmacokinetics of rhBMPs. One of the critical issues is the relationship between the rhBMP2 pharmacokinetics and the osteoinductive activity. The fact that rhBMP-2 exerts its osteoinductive activity in a dose-dependent manner [116] is indicative of the direct effect of at least one of the

pharmacokinetic parameters (initial burst release, local concentration or the slow release rate) on the osteoinductive activity. However, dose-ranging studies do not allow us to pinpoint the critical variable since all of the pharmacokinetic variables are affected with a dose change. To determine the critical variable, carrier–rhBMP combinations with distinct pharmacokinetics at a given implantation dose had to be employed. For this end, a truncated isoform of rhBMP-2 was prepared by cleaving it with plasmin in vitro and purifying the plasmincleaved rhBMP-2 by CEX [117]. Plasmin cleavage removes a positively charged fraction of N-terminus giving a protein with lower pI. This N-terminus region is believed to be involved in heparin and / or other proteoglycan binding [118] but not in osteoinductive activity, being outside the critical 7-cysteine region of the protein. The plasmin-cleaved protein is active in vitro in a W-20 cell assay [119], exhibiting greater activity than the native protein, possibly due to lack of non-specific interactions with extracellular matrix molecules [120]. The pharmacokinetics experiments (in the rat ectopic assay) indicated a rapid burst release of plasmin-cleaved rhBMP-2 from the collagen sponges ( . 80% release within 3 h as opposed to | 30% release by native protein), followed by a similar rate of loss for both types of rhBMP-2. Because of the initial difference, however, the concentration of native rhBMP-2 was | 3–5 times more than the plasmin-cleaved rhBMP-2 at all times. The osteoinductive activity of the plasmincleaved protein was significantly less than the native rhBMP-2 when implanted at equal doses. The plasmin-cleaved rhBMP-2 was able to give robust bone formation at relatively high implant doses. The initial rhBMP-2 burst, therefore, did not seem sufficient for a strong osteoinductive activity and an association between local rhBMP-2 retention and the osteoinductive activity was observed. It is not possible to completely rule out the importance of initial burst (since both systems exhibited burst effects) but local concentration of the osteoinductive factor seemed more critical for the overall osteoinductive activity.

3.3. Alternative carriers An important consideration for the use of osteoconductive scaffolds as BMP carriers concerns the

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Fig. 2. rhBMP-2 pharmacokinetics in the rat ectopic model. (A) rhBMP-2 was implanted with either a reconstituted collagen sponge or DBM of 1–3 mm particle size (in a gelatin capsule). The initial rhBMP-2 retention was different between the two carriers but not the subsequent rhBMP-2 loss. (B) rhBMP-2 pharmacokinetics from sponges with different chemical treatments (formaldehyde crosslinking and ethylene oxide sterilization). The uncrosslinked and unsterile sponges had the highest initial retention but the fastest rhBMP-2 loss. Sterilization did not affect both the initial retention and the subsequent protein loss. (C) Effect of succinylation on rhBMP-2 pharmacokinetics. Both the initial retention and the subsequent release was faster for the chemically modified protein. The dose retained at the implant site (given by pAUC) was | 100-fold less for succinylated protein.

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volume of the defect to be repaired. Osteoconduction is feasible on relatively small osseous defects (regeneration capacity of bone), whereas osteoinductive implants are used or designed for larger size defects where spontaneous bone growth is not feasible. This necessitates a larger volume of biomaterial implantation. It is imperative that the relatively larger volume carriers are resorbed in a suitable time frame to minimize the volume of fibrous tissue and not to interfere with the remodeling of the deposited bone. This point was illustrated in a canine supra-alveolar periodontal defect model by Sigurdsson et al. [68] who compared the efficacy of various carriers (DBM, bovine bone mineral, a collagen sponge, PLGA and PLA granules) for rhBMP-2 implantation. Although the chosen carriers were different in physicochemical nature, substantial bone and cementum regeneration was evident with all carriers. Some of the carriers (bone mineral and PLA granules), however, did not degrade significantly after 8 weeks, thus limiting the amount of newly formed bone. The large volume of the defect also hampered the cohesivity of the carrier, especially for particulate carriers which appeared to migrate from the implant site. Sequestering agents, such as blood [67] and or water-soluble polymers [121], may need to be employed for reproducible osteoinduction with particulate BMP carriers. The carrier geometry has also been proposed to significantly affect the osteoinductive activity of implanted BMP [54,122]. In the rat ectopic assay, significant osteoinductive activity was observed when osteogenin (i.e., BMP-3) was implanted with a disc-shaped hydroxyapatite (HA) carrier, but not with a granular carrier. With the disc-shaped carrier, bone induction was mostly intramembranous with distinct islands of hypertrophic chondrocytes, in contrast to a DBM / BMP combination, where endochondral bone formation is typical. In contrast to DBM, where abundant chondrogenesis was obtained in the rat ectopic model, Kuboki et al. [122] observed direct bone deposition with a porous HA carrier. Similar to the DBM observations, a fibrous glass matrix was shown to induce predominantly chondrogenic tissue within the pores of the implant, while mineralized tissue appeared only at the periphery of the carrier [122]. The authors speculated that small pore size, sufficient for invasion of

mesenchymal stem cells but not for angiogenesis, is likely to inhibit the mineralization of the chondrogenic tissue in fibrous glass matrix. These observations indicated that the geometry of the carrier was able to modulate the osteogenic activity of BMPs and easy vascular access was paramount for the rapid induction of mineralized tissue. Control of particle geometry is feasible for rigid carriers (such as mineral or polymer-based), but more challenging for carriers of low mechanical strength (such as DBM and reconstituted collagen sponges). The observed differences in osteoinductive activity based on carrier geometry raises additional design considerations for an ideal BMP carrier. The protein affinity of BMP carriers is an important issue which is not relevant for osteoconductive carriers. In the rat ectopic model, significant differences in the rhBMP-2 pharmacokinetics were seen among the commonly used osteoconductive biomaterials (e.g., OsteoGraft TM , ProOsteon TM ). Depending on the carrier, we observed a rhBMP-2 loss of 25 to 90% of the implanted dose after 3 h of implantation [117]. Losing 90% of the implanted dose in the first hours of implantation is certainly not desirable. The subsequent release kinetics was also variable with bi-exponential t 1 / 2 ranging from 2 to 6 days [117], indicating significant variations in the affinity of rhBMP-2 for various carriers in vivo. Unlike collagenous carriers, mineral-based carriers appeared to bind a fraction of rhBMP-2 (typically 5–10% of implanted dose) irreversibly (i.e., without release). The biological consequence of the tightlybound BMP fraction is currently not known. A fruitful avenue of research could be the development of composite carriers in which beneficial aspects of more than one biomaterial was combined [66]. Examples of such biomaterials include collagen–proteoglycan scaffolds [123], mineral particles in a collagen matrix (such as Collagraft TM [124]), mineralized collagen sponges [69] and PLA carriers infiltrated with hyaluronic acid [90]. Gunesekeran et al. [69] assessed the osteoinductive activity of a BMP in a rabbit skull defect, comparing the performance of a mineralized collagen carrier to that of collagen, fibrin glue, PLGA and autologous bone as the BMP carriers. When implanted alone, only the mineralized carrier gave osteoconductive activity comparable to that of autologous bone (which is also

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osteoinductive). Collagen, fibrin glue and PLGA were not effective in delivering BMP-2, but mineralized collagen readily gave bony bridging when implanted with BMP-2. Although the reasons for the success of mineralized collagen were not proposed, it is likely that the stable porosity exhibited in vivo, the mechanical properties (due to mineral component), and better cell compatibility (due to collagen component) might have been optimal for osteoinduction. Understanding the mechanisms responsible for success will require continued refinement and, in combination with scaffold technology, undoubtedly will provide a foundation for the development of improved tissue engineering strategies.

4. Tissue engineering strategies Tissue engineering is an emerging field in science that has great potential in clinical medicine [125]. As the name implies, tissue engineering generally utilizes matrix or scaffolding technology, which can be natural, synthetic, or a combination of both, that is interactive with living cells to organize into a threedimensional functional living tissue [125]. Thus, tissue engineering combines elements from the fields of cell transplantation and biomaterials that are intended to provide replacement of diseased tissues / organs in which there is currently a severe shortage. In the following section, we have briefly attempted to summarize some of the emerging technologies that will impact the development of tissue engineering and its applications in the musculoskeletal systems of people with debilitating conditions. One of the exciting areas of bone tissue engineering is the development of ‘smart biomaterials’ [71]. This concept relies on the ability of an implanted biomaterial to elicit a desired morphogenetic response from the host based on the geometry and chemical nature of the material [71]. Additionally, it should be possible to control porosity, biodegradation and BMP affinity of a carrier independent of one another. This control will be necessary for different anatomical sites in which the osteoconductive processes might differ (e.g., in a fracture site vs. an intramuscular spinal fusion site). The rate of degradation is important, since it can provide the balance between in vivo retention of the biological factor vs.

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physical hinderance for the deposited mineralized tissue. The extracellular matrix proteins can obviously modulate the osteoinductive activity of the BMPs and the composition of the deposited mineralized tissue [79]. Recent evidence from cell culture systems have suggested that rhBMPs are able to induce some of these proteins in cell lines and in osteoprogenitor cells [119,126]. Supplying these extracellular matrix proteins in the form of bioactive scaffolds along with osteoinductive factors such as rhBMPs may help to optimize the osteoinduction process. In addition, BMPs alone, or in concert, have been shown to induce total regeneration of bony tissue [71]. Optimizing therapies for osseous repair may involve several BMPs delivered by appropriate carriers / scaffolds. In addition, other growth factors may act in concert with BMPs and potentiate the pharmacological activity of the BMPs / rhBMPs. Once the inter-relationships between different growth factors are characterized, and defined in a systematic reproducible system, carriers will be needed that can control the delivery of multiple growth factors in an orchestrated fashion. The ultimate delivery vehicle for osseous regeneration may be developed in the very near future. The generation of human mesenchymal stem cells as a source to provide the embedding of osteogenic precursors is steadily becoming more feasible [127]. Issues surrounding immunologic incompatibility of allogeneic-sourced tissues may limit this technology, however, a small bone marrow biopsy from a patient and subsequent expansion and seeding onto optimal carriers / scaffolds is feasible. In addition, evolving technologies are under investigation to generate a universal osteogenic cell for bone tissue engineering. Lastly, genetic engineering of a universal donor cell with rhBMP-2, or other factors, may provide delivery of these augmentation molecules in a local, sustained manner at physiologic levels without the need for carrier / factor engineering.

5. Clinical applications Speculations as to the safety and efficacy concerning the clinical application of BMPs will continue until a satisfactory level of data is generated.

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Clearly, the concern is legitimate, based on the limited number of reports linking osteosarcomas with BMP activity [128–130]. However, we have not observed any evidence to date that would lend support to BMPs role as an oncogenic factor. Indeed, BMPs are factors that appear to play a greater role in recruitment and differentiation than factors promoting malignancy. Historically, crude native human BMP preparations had been previously utilized in several institutional review-based studies [131–133]. Urist and Dawson used chemosterilized autolyzed antigen-extracted allogeneic (AAA) bone implants instead of autologous bone grafts for posterior lumbar intertransverse process fusion [131]. The control group was similar patients receiving autogenous bone grafts. The results were similar between the two groups, while the authors stressed that the AAA recipients avoided the morbidity associated with the autologous bone graft harvesting. Additionally, Johnson et al. reported their results in an uncontrolled study of implanting crude native human BMP preparations in refractory tibial and femoral nonunions [132,133]. In patients that experienced unsuccessful multiple surgical grafting, with obvious nonunions, patients receiving the BMP supplementation exhibited osseous union. These studies clearly utilized preparations that were impure, heterogeneous and likely their nature was of a low dose. Nevertheless, these results were suggestive of the safety of BMP preparations in osseous regeneration. Collagen sponges are currently being evaluated in a preliminary clinical study as a sustained release system for rhBMP-2 for maxillary sinus floor augmentation [44]. In this study, the maxillary alveolar sinuses in 12 patients were treated to promote bone formation suitable for dental implants. The observations made after 16 weeks suggested that this combination of rhBMP-2 with absorbable collagen sponge was efficacious and therefore may be an acceptable alternative to traditional treatments for maxillary sinus floor augmentations. In addition, other FDA approved trials have been reported to be underway with tibial fractures and avascular necrosis [70]. However, more studies, especially incorporating other carriers and long-term data (greater than 1 year), must be generated prior to assessing the treatments as clinically acceptable.

6. Conclusions The ability to engineer carrier systems to provide sustained release of the specific bone formation substance known as the bone morphogenetic proteins (BMPs) is fast approaching realization as a practical and functional alternative to conventional therapies. This review has presented the candidate materials poly(a-hydroxy acids) and collagens as exhibiting the potential for safe and well characterized carrier systems. In addition, the clinical therapeutic role of human recombinant BMP-2 delivered from these sustained release carrier systems has been discussed for their ability to provide efficient therapies for the repair of critical-sized osseous defects. However, additional preclinical studies are necessary to determine regional carrier requirements for structure, release kinetics, dose–response and, most importantly, assure the safety of the scaffolds by characterizing the long-term host response prior to undertaking clinical applications. Nevertheless, sustained release systems for BMP delivery hold great promise as a clinical therapeutic.

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