Emerging biomaterials and tissue engineering [1 ed.]

Table of contents :
Preface
Bone morphogenetic proteins: A review for cranial and maxillofacial surgery
Guided bone regeneration
In situ forming biomaterials
Surface modifications of implants
Recent advances in cosmetic materials
Skin and oral mucosal substitutes
Materials and techniques in maxillofacial prosthodontic rehabilitation
Wound closure materials
Cartilage regeneration
Peripheral sensory nerve regeneration with biodegradable materials and neurotropic factor
Index

Citation preview

Oral Maxillofacial Surg Clin N Am 14 (2002) ix

Preface

Emerging biomaterials and tissue engineering

Michael J. Buckley, DMD, MS John C. Keller, PhD Guest Editors

In the future, the practice of oral and maxillofacial surgery will be shaped by the tools we are given. The emerging fields of tissue engineering and biomaterials are truly exciting. Products of the future, including those from genetically engineered proteins all the way to tissue-engineered organs, are and will be in clinical trials in the very near future. Our specialty will benefit from the development of many of these biomaterials, and our future will be shaped by their use. Our patients will benefit from us understanding and critically evaluating new products as they are brought to market. This issue contains articles authored by a diverse group of scientists and clinicians representing a variety of fields. The articles have all been written by authors with great insight into their particular topic. They are written to bring to light the present state of the art, with emphasis on future development of products in that arena. The use of

emerging biomaterials and tissue-engineered approaches to surgical treatment is exciting and will help the advance of oral and maxillofacial surgery. The future is bright.

Michael J. Buckley, DMD, MS University of Pittsburg School of Dental Medicine 3501 Terrace Street Suite 440, Salk Hall Pittsburgh, PA 15213, USA John C. Keller, PhD Dows Institute for Dental Research Dental Research, NY18, DSB University of Iowa Iowa City, IA 52242, USA

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Oral Maxillofacial Surg Clin N Am 14 (2002) 1 – 14

Bone morphogenetic proteins A review for cranial and maxillofacial surgery Kodi Azari, MDa, John S. Doctor, PhDb,c, Bruce A. Doll, DDS, PhDb,d, Jeffrey O. Hollinger, DDS, PhDb,* a Division of Plastic Surgery, University of Pittsburgh Medical Center, Pittsburgh, PA, USA Bone Tissue Engineering Center, Carnegie Mellon University, 125 Smith Hall, 5000 Forbes Avenue, Pittsburgh, PA 15213, USA c Department of Biological Sciences, Duquesne University, Pittsburgh, PA, USA d School of Dental Medicine, University of Pittsburgh, Pittsburgh, PA, USA

b

Along with vitamin D and the Iliazov callous distraction technique, the discovery of bone morphogenetic proteins (BMPs) is deemed one of the three greatest scientific advances in the domain of bone research of the twentieth century [105]. To date, a Medline search of the available literature yields more than 2100 publications that focus in some manner on BMPs. In recent years the use of recombinant DNA technology has led to the identification and molecular cloning of bone morphogenetic proteins resulting in the production of quantities suitable for clinical applications [75,76,110 – 112]. This article serves as a review of the history, mechanisms of signaling, and synopsis of the experimental and clinically relevant information available in the field of BMPs as it pertains to cranial and maxillofacial reconstructive surgery.

Historical perspectives In the late 1800s, Senn [93] implanted demineralized ox bone and iodoform to treat dog skull osteomyelitis and noted better than expected bony defect healing. Later, Lexer [51] suggested that necrotic

John S. Doctor was supported by the National Institutes of Health sabbatical fellowship F33 AR08268. * Corresponding author. E-mail address: [email protected] (J.O. Hollinger).

bone tissue released stimulating factors that affect osteoblasts. By the 1920s it was hypothesized that a substance released from graft tissue resulted in the differentiation of fibroblasts into bone and cartilage forming cells [79]. The next hypothesis, proposed in 1928, was that calcium materials contained in the graft tissue were the putative agents that induced new bone formation [50]. As a means of managing visceral defects, however, Neuhof [72] implanted fascia along the urinary tract and noticed local bone formation. This finding suggested that extraskeletal ossification may be accomplished by means other than transplanting calcium-containing grafts. In some pathologic conditions, extraskeletal tissues may ossify and exhibit all the features of bone, including marrow [22]. In 1903, Binnie [8] described myositis ossificans traumatica, in which frequently repeated soft tissue trauma resulted in heterotopic ossification of muscle. Other causes of ectopic bone formation include tumors of chondrocytic and osteocytic origin and the uncommon connective tissue disease called myositis ossificans progressiva, which, as the name implies, leads to progressive soft tissue ossification and bony skeleton malformations [8,22]. In 1965, in an experiment to recalcify cortical bone, Marshall Urist [99] made the key discovery that led to the hunt for factors responsible for bone formation. In this classic experiment, demineralized bone fragments were implanted subcutaneously or in intramuscular pouches in rats and rabbits. Several weeks after implantation new cartilage and bone appeared at the implantation sites. This occurrence led to the hypoth-

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esis of bone formation by autoinduction, in which grafted inducer cells influence resident primitive cells to differentiate into osteoprogenitor or chondroprogenitor cells. Urist’s discovery led to a series of investigations to determine the putative agents in demineralized bone that provoke autoinduction. The osteoinductive activity in bone matrix was found to be the result of a noncollagenous and water soluble substance coined BMP [100 – 102,104] The next major advance was the purification of BMPs and the isolation of 16, 18 and 30 Kda polypeptides from bovine bone. Later, these proteins were named BMP 1-3 [106]. In the late 1980s, recombinant DNA technology was used to generate human complementary DNA clones of BMP molecules [112]. These clones were inserted into mammalian and nonmammalian cells to produce recombinant BMPs. Thus far more than 15 BMPs have been isolated and reported in the literature [74].

Mechanisms of BMP signaling The BMPs are members of the transforming growth factor-b (TGF-b) family of growth and differentiation factors. Although identified and named because of their osteoinductive activity [14,110], the BMPs play diverse roles during embryonic and postembryonic development as signaling molecules in a wide range of tissues [30,31,40,94]. For example, in addition to roles in skeletogenesis and bone formation [83], BMPs are implicated in mesoderm patterning, neurogenesis, and organogenesis [62,64,66,77, 84,109]. In the last several years, genetic, molecular, and cellular approaches have provided important details about the mechanisms of BMP signaling. Much of our knowledge derives from studies on model organisms, including mouse, frog (Xenopus), nematode (Caenhorabditis), and fruitfly (Drosophila). As indicated schematically in Fig. 1, extracellular BMPs bind and activate a multimeric transmembrane receptor complex. This ligand-activated receptor, serine/ threonine kinase, phosphorylates members of the Smad family of signal transducers, which results in the direct translocation of Smads to the nucleus, where they can modulate the expression of target genes [58,113]. This description only hints at the complexity inherent in the control of BMP signaling. If BMPs are to serve as effective therapeutic agents [87,90], it is necessary to understand their action, including (1) their role in signaling a multitude of cell types, (2) mechanisms by which BMP signaling can be modulated

extracellularly and intracellularly, and (3) the potential for agonistic and antagonistic activities of BMPs depending on the particular cell type. The amount of crosstalk between the BMP signaling pathway and other growth factor and hormonal signaling pathways [46,48,65] must be evaluated as to avoid unwanted side effects from BMP-based therapies. The BMPs are members of a large family of signaling proteins that are important during development and share the same basic structure The TGF-b superfamily contains several dozen members, including the BMPs, the activins, mu¨llerian inhibiting substance, and the prototypic TGF-bs. More than a dozen proteins are grouped into the BMP family, based on sequence similarity. The roles of BMPs during development can be inferred from their expression in various tissues [28,31,54]. Genetic analyses in mice demonstrate roles for BMPs in skeletogenesis and other aspects of development. For example, several mutations that perturb the function of BMP genes are known, including, short ear and brachypod. Short ear, which results in abnormal growth and patterning of skeletal structures and diminished repair of bone fractures in adults, is caused by lesions in the BMP-5 gene [44,45,57]. Mutations in GDF-5 (growth and differentiation factor-5, a member of the BMP family) result in the brachypod phenotype in mice and in the autosomal recessive syndromes Hunter-Thompson and Grebe-type chondrodysplasias in humans [24, 95]. These syndromes are characterized by the shortening of the appendicular skeleton and loss or abnormal development of some joints, with little effect on the axial skeleton. Directed disruptions of BMP genes in mice reveal important functions in mesoderm induction and organogenesis [20,53,109] but have been less informative about the roles of BMPs during skeletogenesis. For example, BMP-2 and BMP-4 knockout mice die early in embryonic development, long before development of the skeleton, because of defects in gastrulation [109,116]. BMP-7 knockout mice exhibit eye and kidney defects but only mild skeletal defects, perhaps because of the rebundant function of other BMPs [53]. Molecular analysis reveals that BMPs share the major hallmarks of all members of the TGF-b superfamily: initially to facilitate secretion, they are synthesized as large precursors that consist of an amino-terminal signal sequence, a pro-domain and a carboxy-terminal bioactive mature domain [13,111, 112]. The amino-terminal signal sequence and prodomain regions of the BMPs vary in size and

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Fig. 1. Signaling by bone morphogenetic proteins (BMPs). The binding of an extracellular BMP dimeric ligand to the multimeric BMP transmembrane receptor complex. Formation of the ligand/receptor complex results in the activation of the serine/threonine kinase of BMP receptor-1 (BMPR-I) and the phosphorylation of members of the R-Smad family of signal transducers (Smads 1, 5, or 8). This phosphorylation can be antagonized by the action of inhibitory Smad 6 (not shown). The phosphorylated R-Smad binds with C-Smad 4 and translocates into the nucleus where R-Smad/C-Smad complex directly or indirectly interacts with the core binding factor (CBF). The activated CBF translocates to the osteocalcin promoter region and activates Osteoblast-Specific Element-2 (OSE-2). This sequence of events results in the transcription of the osteocalcin gene. Osteocalcin mRNA is translated to osteocalcin protein on the ribosomes. P = phosphate. mRNA = messenger ribonucleic acid (Adapted from Schmitt J. Bone morphogenetic proteins: an update on basic biology and clinical relevance. J Ortho Res 1999;17:269 – 78; with permission).

sequence, whereas the mature domain shows greater sequence similarity among family members (Fig. 2). The mature domain is usually cleaved from the prodomain region by the subtilsin-like serine protease furin [6,18] to produce an active polypeptide of between 110 and 140 amino acids in length. The mature region of BMPs is highly conserved among the family members and includes seven nearly invariant cysteines. Six of these are involved in the formation of intrachain disulfide bonds to make a rigid structure called a ‘‘cysteine knot’’; the seventh is involved in the formation of homodimers and heterodimers by an interchain disulfide bond [89].

During intracellular processing before proteolytic cleavage and secretion, these mature domains form either homodimers made up of two of the same family members or heterodimers made up of two different family members to produce the active signaling molecule. Depending on the expression of BMPs in a particular cell type, it is possible to form combinations of various homodimers and heterodimers, each with potentially differing activities. Overlapping expression of murine BMPs supports this contention for the combinatorial effects of homodimeric and heterodimeric BMPs [54]. Heterodimers of Xenopus BMP-4 and BMP-7 produced in vitro [27] are more potent

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Fig. 2. The polypeptide structure of BMPs.

than homodimers in bone- and mesoderm-inducing activities [2,96]. The pro-domain seems to be required for normal synthesis and secretion of TGF-b superfamily members, and it can remain associated with the mature domain to produce an inactive latent complex for some members of the family [86] but apparently not for the BMPs. Other proteins can associate with extracellular BMPs with important implications for their biologic activity, as described in the following section. Several extracellular proteins act to regulate the activity of BMPs The extracellular environment seems to play an important role in regulating the activity of BMPs [25]. In addition to interactions with extracellular matrix components such as collagens and heparin sulfate proteoglycans, BMPs interact specifically with several extracellular proteins. The protein noggin binds BMP-2 and BMP-4 with highly affinity and inhibits their association with BMP receptors (BMPR) [119]; noggin also binds BMP-7 with low affinity. Another secreted protein, follistatin, originally identified as a high affinity antagonist of activin [68], also has affinity for BMP-4 and BMP-7 [23,37]. Studies of Xenopus embryos also implicate the DAN family (cerebus, gremlin, DAN) as antagonists of BMP signaling by direct binding to BMPs, thereby blocking ligand/receptor interactions [35]. Similar to noggin, the protein chordin binds BMP-2 and BMP-4 with high affinity and inhibits BMP/receptor interaction [78]. In the case of chordin, members of the astacin family of metalloproteases (including BMP-1 and several tolloid-like proteases) cleave chordin, thereby releasing BMPs to interact with BMPRs [91]. Components of the extracellular environment act to sequester or release BMPs (Fig. 3). Chordin/BMP complexes may serve as an extracellular reservoir for BMPs, with the level of the reservoir depending on the amounts of chordin, BMP, and chordin-specific BMP-1/tolloid proteases. Based on observations in rat pituitary cells that follistatin/activin complexes are

internalized and degraded, follistatin may act to clear activin and perhaps, BMPs from the extracellular environment. Attention to these extracellular regulators is important in designing therapies to deliver and release BMPs. BMP receptors are heteromeric complexes with serine/threonine kinase activity Bone morphogenetic protein members interact with the extracellular domain of a family of cell surface type I and type II receptors to signal across the cell membrane and elicit a cellular response [41,58]. A complex of a type I receptor and a type II receptor assembles through interaction with a BMP ligand. Formation of this ligand-receptor complex results in the phosphorylation of the type I receptor’s kinase domain by the serine/threonine kinase domain of the type II receptor. This activation of the type I receptor’s serine/ threonine kinase results in the downstream phosphorylation of target substrates, including the Smad family of signal transducing proteins (see Fig. 1). In mammals, BMP type I receptors (BMPR-IA and BMPR-IB) and BMP type II receptors (BMPRII) are expressed during embryogenesis and in several cell lines [19,21,28,71]. Evidence for important roles of BMPRs during development comes from analysis of defective BMPR-I receptors (loss-of-function experiments) [71,97] that block BMP signaling and constitutively active receptors (gain-of-function experiments) [1] that elicit characteristics cell responses in the absence of BMPs. Important for a role in osteogenesis, BMPRs are expressed during bone formation in skeletal development and in fracture repair [19,38]. Mutations in the mouse brachydactyly gene, which encodes BMPR-1B, indicate a role of BMPRs in skeletogenesis [7]. Smads are essential components of intracellular signaling by BMPs The discovery of Smads, the intracellular downstream targets of BMPR-II kinase activity in mam-

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Fig. 3. BMPs interact specifically with extracellular proteins that act as antagonists to sequester BMPs from signaling. The inhibitory activities of distinct types of proteins on extracellular BMPs are depicted. Noggin follistatin, chordin, and members of the differential screening-selected gene abberative in neuroblastoma (DAN) family (cerebus, gremlin, DAN) bind to members of the TGF-b family, including BMPs, and inhibit their ability to bind cognate receptors. The antagonistic activity of chordin is itself inhibited by BMP1/ tolloid proteases, which act to degrade chordin, thereby releasing BMPs.

mals, derives directly from genetic analysis of mutations that affect the signaling activity of the fruitfly BMP-2 and -4-like protein known as decapentaplegic. Work in Drosophila on the Mad gene (Mothers against decapentaplegic), which mediates decapentaplegic signaling [92] and on a series of sma genes in the nematode Caenorhabditis [88], led to the isolation of sma/Mad-like genes, termed the Smads in mammals [4]. To date, genes that encode eight related mammalian Smad proteins, Smad1 through Smad8, have been identified and characterized [16,29,41,113]. These proteins are grouped into three categories: the receptor-regulated, pathway-restricted R-Smads (Smad1, 2, 3, 5, 8); the common mediator C-Smad (Smad 4); and the inhibitory I-Smad (Smad5, 6). The R-Smads are activated by BMPR-I phosphorylation of three key serines in their effector domain in a ligandspecific fashion. Smad1, Smad5 and Smad8 are involved in BMP signaling, whereas Smad2 and Smad3 mediate TGF-b/activin signaling [4]. Smad4, the C-Smad, is not a substrate for receptor kinases. Instead, it forms a heterodimeric complex with a phosphorylated R-Smad, and this complex translocates to the nucleus to modulate the expression of target genes (see Fig. 1). R-Smads contain a motif essential for translocation of the R-Smad /C-Smad complex into the nucleus [114,117]. The I-Smads inhibit signaling by binding to the intracellular domain of the receptor. For example, Smad6 binds in a stable fashion

to BMPR-I and competes for binding by R-Smads [39]. Smad6 also can inhibit BMP signaling by competing for the binding of phosphorylated R-Smads to Smad4 [26]. R-Smad/C-Smad complexes interact with transcription factors in the nucleus to modulate gene expression Once in the nucleus, R-Smad/C-Smad complexes can bind directly, although with low affinity, to DNA in the regulatory regions of many BMP-responsive genes [47,59]. The formation of stable DNA-binding complexes requires the interaction of Smad complexes with sequence-specific DNA binding proteins [41]. Several such DNA binding proteins have been identified for acitvin-specific Smads2 and 3, including FAST-1, FAST-2, and AP-1 [15,49,116]. In the case of BMP Smads1, 5, and 8, several recent papers report Smad interactions with SIP1, the acute myelogenous leukemia protein, Hoxc-8, a homeodomain containing transcription factor, and p300/CBP and STAT3, transcription factors important in astrocyte differentiation [59,70]. A particularly interesting result implicates a Smad interaction with RUNX2/CBFA1, a component of a transcription factor that is essential for proper bone development [118]. Gene knockouts of the RUNX2 gene in mice result in the absence of mature osteoblasts and the absence of bone formation. Muta-

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tions of the RUNX2 gene in humans are implicated in cleidocranial dysplasia, a dominant human bone disease characterized by multiple skeletal abnormalities. A truncated RUNX2 protein from a patient with cleidocranial dysplasia fails to interact with and respond to Smads, thereby implicating impaired Smad signaling in this genetic disease.

Role of BMP after bone injury When the bone, a living tissue, is injured, such as by fracture (Fig. 4), an immediate inflammatory response is elicited. At the injured site vascular endothelial damage results in the activation of the complement cascade, platelet aggregation, and release of its alpha granule contents. This platelet degranulation releases growth factors and triggers chemotactic signals. Polymorphonuclear leukocytes (PMNs), lymphocytes, blood monocytes and tissue macrophages are lured to the wound bed and are activated to release basic fibroblast growth factor (bFGF). The extravasated blood collection clots. Hematoma formation serves as a hemostatic plug to limit further hemorrhage. The conductors of the cascade are the platelets, which have the duty of hemostasis and mediator signaling through the elaboration of chemoat-

tractant growth factors such as platelet-derived growth factor (PDGF), TGF-b, fibroblast growth factor (FGF), endothelial growth factor (EGF), insulin-like growth factor-1 (IGF-1), and platelet factor-4 (PF-4) [3,90]. The early fracture milieu is characteristically an hypoxic and acidic environment that is optimal for the activities of PMNs and tissue macrophages [32]. PMNs phagocytize bacteria and microdebris, whereas larger sized materials are removed by macrophages that may develop into multinucleated giant cells. Macrophages perform myriad functions in the local wound site, including synthesis of growth factors that affect local cell activity, cell recruitment, chemoattraction, and mitogenesis. By days 3 to 5 after bone injury, a local environment develops that includes new capillary sprouts, fibroblasts, macrophages, and collagen isotypes. It is believed that the attachment of growth factors to collagens serves to localize, protect, and arrange these factors for cellular interaction [80]. The collagenous portion of the healing wound renders a key instructional framework to position cytokines and BMPs to receptive cells. Primitive undifferentiated systemically circulating cells, along with osteoprogenitor cells of the periosteum and endosteum, attach to the granulation tissue collagen, which present signaling molecules, including BMP. The cell-ligand interaction

Fig. 4. Sequence of events following bone fracture. Following osseous fracture, platelets degranulate, a hematoma is formed, the immune system is activated at the wound site, and microdebris is removed by phagocytizing cells. By day 3 – 5 postinjury, neovascularization occurs while osteoprogenitor and mesenchymal stem cells localized at the wound site respond to environmental factors, including BMPs, fibroblast growth factor (FGF), transforming growth factor b, insulin-like growth factors (IGF), platelet-derived growth factor (PDGF). (Adapted from Schmitt J. Bone morphogenetic proteins: an update on basic biology and clinical relevance. J Ortho Res 1999;17:269 – 78; with permission.)

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prompts differentiation of these cells into osteoblasts and chondrocytes [80,81]. The influence of BMPs on cell differentiation is of particular interest with respect to bone formation. The combinatorial functions of cell attachment, transduction, and cell-ligand interaction prompt cellular differentiation into specific phenotypes to repair osseous wounds. Cellular armies in the appropriate milieu of growth and differentiating factors such as TGF-b, FGF, VEGF, PDGF, and BMPs eventuate in fracture repair by 6 to 8 weeks after injury [32]. Interestingly, the repaired bone structure is indistinguishable from the preinjury structure. The remodeling of injured bone to the preinjury structure is crafted carefully by osteoblasts and osteoclasts [103]. If, however, a sufficient number of cells are not present at the osseous injury site, they must be recruited, stimulated to proliferate, and acted upon by the appropriate admixture of growth factors. At the fracture site, fragments of the ubiquitous attachment factor fibronectin and degradation products from the extracellular matrix stimulate the transformation of monocytes of osteoclasts [32]. Macrophages at the local wound site elaborate bFGF and VEGF, thus inducing neovascularization to furnish transport pathways for additional cells to replenish those lost to injury [36]. The clinical relevance of BMPs in the regeneration and repair of injured bone requires a responsive regional cell population at the wound site [82]. Cells must be capable of responding to the BMP signal or signals, and sufficient quantity and forms of biologically active BMPs must be attendant to produce the sought outcome, (e.g., to regenerate the form and function of bone). BMPs and their receptors are the executives in this remarkable process [87,90]. Recent studies reveal increased expression of BMP-2, -4 and -7 in the primitive mesenchymal and osteoprogenitor cells, fibroblasts, and proliferating chondrocytes present at the fracture site [9,69,73]. In a rat model, mesenchymal cells that had migrated into the fracture gap and had begun to proliferate showed increased expression of BMP-2 and -4 [9]. In a similar rat fracture healing model, Onishi et al. [73] confirmed that BMP-2, -4, and -7 were present in newly formed trabecular bone and multinucleated osteoclastlike cells. In conclusion, several findings suggest that BMP -2, -4, and -7 work to promote fracture healing and bone regeneration [87].

Bone morphogenetic protein application to correct craniofacial defects Composite grafts that contain purified native BMP were applied in skeletal defects by Sailer [85]. BMP

7

was either added as a paste of purified BMP and granular bone matrix or it was added to lyophilized cartilage or bone during rehydration. BMP paste probably consisted of several homologous proteins [14,112]. After rigorous testing for bone induction in a mouse assay, Muthukumaran et al. [67] purified bovine BMP and delivered it to human skeletal defects. Sailer summarized the maxillofacial reconstruction procedures using bovine BMP between 1992 and 1995 [52]. Age, sex, surgical site, and preexisting traumatic conditions did not hinder engraftment as judged by postsurgical clinical and radiographic examination in the nine cases presented [52]. The use of native, purified BMP is limited by the available source, arduous preparation, and low yields. Small yields of BMP are obtained from kilogram quantities of dry bovine bone. The preparation includes defatting, demineralization, and chromatographic steps—a time-consuming process for yields under 1:50,000. Relative ratios of BMPs within the preparation vary and confound assignment of osteogenic potential for each BMP. Despite the apparent usefulness of cross-species applications, concern for antigenicity encourages use of human recombinant BMPs. The availability of large amounts of highly purified, efficiently produced, predictably constituted recombinant proteins facilitates incorporation into tissue-engineered vehicles [107]. What follows is the description of several different biodegradable composite bone graft materials that are in clinical or preclinical use for the treatment of skeletal defects in orthopedic, maxillofacial, and periodontal surgery.

Animal studies The posterior maxilla is a challenging area in which to place dental implants successfully because of poor bone quality, poor vascularity, and close approximation to the maxillary sinus. Remaining bone often necessitates sinus augmentation. Prior attempts to increase the bone thickness used xenograft, allograft, and autograft materials. These materials have provided therapeutic success but did not result in a consistent and predictable sinus augmentation in all applications. McAllister et al. [61] examined three doses of recombinant human osteogenic protein-1 (rhOP-1; also known as BMP-7), 0.25, 0.6, and 2.5 mg OP-1 per gram of collagen matrix, natural bone mineral, or collagen matrix alone (control) placed in the maxillary sinus of adult chimpanzees. Based on histologic, tomographic, and clinical examination, rhOP-1 treatment groups had greater amounts of bone formation relative to the control group at 7.5 months. Bone

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mineral density and the rate of mineralization were enhanced in the treatment groups. In a related study [56], Margolin et al. demonstrated short-term histologic and radiographic bone healing after sinus grafting with anorganic bovine bone mineral (Bio-Oss) in the chimpanzee sinus. At 18 months, radiographic and histologic evaluation of the chimpanzee sinus grafts demonstrated maintenance of radiographic bone height and bone mineral density [60]. Anorganic bone was replaced by vital bone. The controls exhibited wide variation in parameters insufficient for implant placement. Direct comparisons of recombinant grafts and allografts were not included in these studies. Using a similar collagen-type I sponge, either 0.2 or 0.8 mg/mL rhBMP-2 was added to bilateral hemimandibulectomies in adult male Macaca fasicularis monkeys. Four months after engraftment, endosseous implants were placed in a two-stage procedure. The implants were uncovered and restored to function after 4 additional months. Ridges were regenerated completely on both doses. Histomorphometry dem-

onstrated favorable trabeculation patterns in the resected areas [10]. Table 1 summarizes other animal experiments of interest to the cranial and maxillofacial reconstructive surgeon.

Human studies using bone morphogenetic protein for maxillofacial reconstruction Human studies using recombinant BMP are limited. Eleven patients were evaluated radiographically for overall bone height changes after rhBMP-2 placement in the maxillary sinus [11]. A collagen sponge (ACS, absorbable collagen hemostatic agent) vehicle was used to deliver the rhBMP-2 (dose range 1.77 – 3.4 mg). A dose to bone formation analysis was not performed. Immunologic evaluation revealed no antibody titers to rhBMP-2 or human type-1 collagen. One patient, however, developed an antibody to bovine type-1 collagen yet remained asymptomatic and yielded ‘‘positive bone formation.’’ The vehicle was considered easy to use when form, cohesiveness,

Table 1 Experimental craniomaxillofacial applications of bone morphogenetic proteins: animal studies Author/year

Factor

Species

Defect

Protocol

Results

Barboza et al 2000 [5]

rhBMP-2

Dog

Class III alveolar

Breitbart et al 1999 [12]

BMP-7

Rabbit

Calvarium

Clinically relevant ridge augmentation when HA was added to rhBMP-2 and absorbable collagen sponge BMP-7 addition enhanced bone repair

Khouri et al 1996 [42]

BMP-3

Rat

Calvarium

Marden et al 1994 [55]

rhBMP-2

Rat

Calvarium

Toriumi et al 1991 [98]

rhBMP-2

Dog

Mandible

Khouri et al 1991 [43]

BMP-3 (osteogenin)

Rat

Alveolar ridge augmentation with rhBMP-2/absorbable collagen sponge with or without HA Rabbit periosteal cells with or without transfected human BMP-7 added to critical-sized defects 1500 rad irradiated calvarial defect covered with muscle flap with or without BMP-3 Critical sized defects were treated with rhBMP-2 in an insoluble collagenous bone matrix (ICBM) or with demineralized matrix (DBM) 3 cm defects were plate stabilized and treated with inactive dog bone matrix carrier with or without rhBMP-2 Thigh adductor muscle flaps were placed inside bivalved silicone rubber molds, injected with osteogenin, and coated with demineralized bone matrix

Abbreviation: HA, hydroxyapatite.

Significantly greater healing with BMP-3 Addition of rhBMP-2 was superior to ICBM and DBM alone

By 10 weeks a stiff, noncompressible, mineralized bone formed across the defects, allowing plate removal and chewing a solid diet Flaps treated with osteogenin were entirely transformed into cancellous bone that matched the exact shape of the mold; successfully generated perfused bone in the shapes of femoral heads and mandibles

Author/year

Sample

Boyne et al 1997 [11]

11 patients rhBMP-2 1.77 – 3.4 mg

Howell et al 1997 [34]

Factor/dose

Controls Defect None

12 patients rhBMP-2, mean dose 0.83 mg None (ridge augmentation) and 0.27 mg (extraction sites) Cochran et al 12 patients Follow-up of study 2 None 2000 [17]

Protocol

Insufficient sinus ridge for implant placement

ACS soaked in rhBMP-2 placed in maxillary sinus, biopsy five patients approximately 6 months after surgery Six ridge augmentations, ACS soaked in rhBMP-2, assessment six extraction + graft at 4 months after graft placement Six grafted extraction sites, four grafted ridge augmentation sites

Implants were placed and restored, assessment at 2 years after graft placement

Abbreviations: ACS, absorbable collaqen hemostatic agent (Heliostat); rhBMP-2, recombinant human bmp-2.

Results Immunologic response unremarkable; three patients had subsequent implant placement, woven bone detected in variable amount Healing uneventful, minimal clinically evident ridge augmentation Follow-up for study 2, histologic analysis of graft site similar surrounding native bone

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Table 2 Maxillofacial applications of bone morphogenetic proteins: human studies

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handling, volume, time, and ease of placement were considered. Adverse effects that exceeded more than 5% of total adverse effects included facial edema, oral erythema, mouth pain, ecchymosis, rhinitis, and sinusitis. Approximately 6 months after the operation, biopsy specimens from five patients exhibited a wide variation in woven bone quantity, undetectable or small populations of active bone cells, and capillaries. No residual collagen matrix was noted in any of the biopsy specimens taken after 19 weeks postoperatively. The adverse reactions were consistent with usual morbidity for the sinus augmentation procedure. After therapy with rhBMP-2, three patients did not meet the criteria for implant placement. The authors suggested that the method of access to the sinus may influence efficacy of rhBMP-2-induced bone formation [11]. The alveolar ridge undergoes resorption after loss of the teeth. Edentulous spaces characteristically have decreased bone height and width—a complication for the placement of implants. A two-center clinical trial evaluated rhBMP-2/ACS enhancement of alveolar ridge after tooth extraction in six patients. Second, augmentation of localized alveolar defects was attempted in six patients. The mean dose for rhBMP-2 was 0.27 mg in the extraction sites and 0.83 mg in the augmentation sites. Two aspects for each procedure were evaluated: safety and osseointegration by tomographic analysis over a 4-month period. The Boyne et al. [11] and Howell et al. [34] studies lacked controls. No biopsies were analyzed in the extraction/graft and ridge augmentation study (Table 2). Both papers reported relative ease in the use of the graft vehicle, however [34]. A standardized method of height measurement also was proposed for assessing augmentation. Probing and palpation of the graft sites provided an additional clinical impression of bone integrity. Ridge augmentation with the graft was not observed. The interpretation of any effect by the graft vehicle was confounded by the expected remodeling in extraction sites and the postsurgical bone loss seen after gingival reflection. Efficacy of rhBMP-2 must consider adequacy of dose, duration of exposure, and physical characteristics of the graft sites (soft tissue compression, alveolar wall integrity and number, primary closure) [115]. As with the human sinus augmentations, no specific adverse reaction was attributed to the graft vehicle. Subsequent endosteal implants in six grafted extraction sites or four ridge augmentations were functional and stable at 2 years after graft placement. Healing was uneventful. Histologic features were consistent with normal bone tissue ‘‘identical to surrounding native bone.’’ The authors

suggested that rhBMP-2/ACS (0.43 mg/mL) can be used safely in tooth extraction sites and in local ridge augmentation procedures [17]. Apparently the varied results with construct, growth factor, and site are promising. A tissue-engineered therapy that uses recombinant BMPs is available yet unpredictable. Manipulation of host factors to predispose the site to bone regeneration should accompany the advancement in material available for delivery to the site. A wide range of materials has been used as vehicles or scaffolds for the delivery of BMPs [33,63]. The pharmacokinetics of rhBMP can be influenced by the implant carrier. The likely mechanisms of rhBMP release include (1) desorption of a scaffold, (2) solubility (slow dissolution of sparingly soluble rhBMP), (3) biodegradation of scaffold, and (4) any combination thereof. Using a heterotopic site, Winn et al. [108] suggested that carrier chemistry is a defining but alterable influence on pharmokinetics of rhBMP release. The success or failure to augment bone regeneration must consider constructs of varying geometry and content to judge thoroughly the efficacy of recombinant therapy.

Summary Progress in the area of BMP research has resulted in the knowledge of the mechanisms of BMP signaling from the extracellular environment to changes in gene expression in the nucleus. BMPs are currently being evaluated in animal experiments and human clinical trials for wide-ranging applications. These advances in BMP research have led to the promise and, to some degree, the reality of their application as an alternative to autograft and allogeneic bone grafts to treat craniofacial osseous deficiencies. Future investigations must determine the appropriate doseresponse relationships and focus on the development of a safe and effective carier system to deliver BMP predictably at the implant site.

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Guided bone regeneration Franci Stavropoulos, DDS*, John C. Nale, James D. Ruskin, DMD, MD, FACS Department of Oral and Maxillofacial Surgery and Diagnostic Sciences, University of Florida College of Dentistry, P.O. Box 100416, Gainesville, FL, USA

Currently, the use of guided bone regeneration in the fields of oral and maxillofacial surgery, periodontics, and plastic and reconstructive surgery is a predictable and effective technique to prevent epithelial and connective tissue migration into an osseous defect. Osseous defects that develop either congenitally or traumatically, as a result of infection, or secondary to the surgical resection of a tumor, challenge the surgeon whose goal is to reconstruct the defect to its original anatomic and functional state. Significant amounts of research have focused on the evaluation of bone grafting materials and surgical techniques; however, the preservation of the grafted site or the promotion of new bone is addressed by the use of a barrier membrane.

Historical development Barrier membranes were first recognized in the orthopedic literature for the healing of bone defects. In 1957, Murray et al. [32] published an article in the American Journal of Surgery discussing the influence soft tissue played on preventing bone from growing into an osseous defect. They stated that only three conditions were necessary for the new growth of bone into a surrounding defect: (1) the presence of a blood clot, (2) preserved osteoblasts, and (3) contact with living tissue. Experimentally, a rigid plastic cage was placed over a small defect created in the ilium of dogs; after 10 weeks of healing, the specimens were recovered. New bone growth was evident under the cage.

* Corresponding author. E-mail address: [email protected] (F. Stavropoulos).

They concluded that excluding soft tissue and creating space that protected the hematoma made complete bone fill possible. The only limitation to the amount of new bone formed was the height of the cage as determined by the overlying soft tissue [32]. Hurley et al. [20] continued experimenting with two types of barrier membranes in spinal fusion procedures in dogs, again supporting the theory that the membrane prevented the overlying differentiated soft tissue from encroaching into the area of newly forming bone. The benefit of Millipore (Millipore Corp, Bedford, MA) as a semipermeable membrane, which isolated bony defects from the surrounding soft tissue in the radii of canines, was reported in 1961 by Bassett et al. [2]. Melcher and Dryer [29] reported in 1962 on the healing of bony defects in the femur of adult rats when the defect was protected by either an intact, cellulose-acetate shield or homogenous organic bone. They concluded that new bone developed in the defects covered by either membrane; however, the maintenance of the shape of the new bone depended on the nonresorbed membrane. It was again recognized that the barrier shield protected the hematoma from invasion of connective tissue and prevented distortion of the hematoma by the overlying soft tissue [29]. In 1968, Boyne and Mikels [5] first introduced the concept of excluding soft tissue with the use of a filter material for bone regeneration to the oral surgery literature. A section of bone was removed from the inferior border of adult, canine mandibles, and the residual space was then fitted with a wire mesh lined with a cellulose-acetate filter. Before placement, bone marrow from the femur of the dog was packed densely into the mesh cellulose-acetate crib. Specimens obtained 12 weeks after surgery revealed proliferating bone beneath the mesh-filter implant.

1042-3699/02/$ – see front matter D 2002, Elsevier Science (USA). All rights reserved. PII: S 1 0 4 2 - 3 6 9 9 ( 0 2 ) 0 0 0 1 3 - 4

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Biopsy specimens also showed new bone growth extending along the entire inner surface of the filter, with fibrous tissue remaining only along the outer surface of the filter [5]. Generalized, accepted clinical application of those pioneering studies did not occur until the 1980s, when the potential of the technique was recognized by Gottlow et al. [17]. They systematically examined the relative contribution of varying cells that, after periodontal therapy, potentially could colonize the adjacent root surface and dictate the quality of the newly formed tissue in the periodontium. For instance, root resorption was found to occur secondary to the presence of granulation tissue. Reepithelialization of the root surface also prevented connective tissue attachment and root resorption. The group found the cells from the periodontal ligament to be the only cells capable of forming a new attachment. The concept of guided tissue regeneration was introduced and subsequently tested [17]. Nyman et al. [33] treated 11 periodontally involved teeth in 10 patients and performed periodontal therapy without osseous surgery. A teflon membrane (Gore-Tex, WL Gore & Associates Inc, Flagstaff, AZ) was placed on the coronal one third of the root. The mucoperiosteal flaps were secured with sutures over the membrane. Three months after healing, four teeth were extracted with their associated periodontal tissues and examined histologically. New attachment was evident in five sections. The seven nonextracted teeth were evaluated for attachment level at a second, surgical procedure to remove the membrane 3 months after placement and again 3 months later, using clinical parameters of probing attachment levels. In some teeth, new attachment formed, whereas in others, only a few millimeters of new cementum formed [33]. These results demonstrated the basis of the biologic principle of guided tissue regeneration for the treatment of periodontal disease in humans. Currently, the periodontal literature is replete with basic research and clinical investigations concerning guided tissue regeneration for the treatment of localized bone defects in natural dentition and in conjunction with endosseous implants.

Guided bone regeneration Based on the previous findings, barrier membranes again were studied experimentally for the regeneration of bone defects in the maxilla and the mandible. Dahlin and colleagues [10 – 13] spearheaded the research on guided bone regeneration in an attempt to solve the confounding problem of re-

constructing large, osseous defects in the jaws and for the treatment of the atrophic maxilla or mandible. Numerous methods have been developed for the augmentation of such defects, including autogenous bone grafting, allografts, and demineralized bone preparations, all with various degrees of success at promoting osteogenesis. It is known that to accomplish the repair of a bone defect, the rate of osteogenesis extending inward from the adjacent bone ends must exceed the rate of fibrogenesis growing in from the surrounding muscle or connective tissue [48]. Guided tissue regeneration was again looked to for assistance in complete osteogenesis. In 1988, Dahlin et al. [12] published the results of animal experimentation on the healing of bone defects. Bilaterally, a through-and-through defect was surgically created in the ramus in 30 SpragueDawley rats. On one side, the defect was covered with a porous (0.45 mm) polytetrafluoroethylene (PTFE) membrance (Gore-Tex). The other side served as the control, without a membrane covering. After 3, 6, and 9 weeks of healing, the specimens were evaluated macroscopically and histologically by light microscope. Statistical analysis of the heated sites demonstrated a highly significant ( P < 0.001) increase in bone regeneration on the membrane side as compared to the control. Therefore, it was demonstrated that soft tissue ingrowth into a bony defect could be prevented and consequently could enhance unimpeded bone healing greatly. Further experimentation was performed in animals which demonstrated the generation of bone around titanium implants. One of the essential criteria for the successful placement and osseointegration of an endosseous implant is a sufficient volume of bone, one that completely covers the threads of the implant. It is known that a minimum of 7 mm  4 mm of bone is required when the smallest type of implant fixture is used in the oral cavity [24]. As an alternative to a two-staged, bone grafting procedure followed by implant placement, Dahlin et al. [13] evaluated the principle of guided tissue regeneration to generate bone at the exposed parts of titanium implants. Thirty ‘‘commercially pure’’ 10-mm titanium implants (Nobelpharma, Gothenburg, Sweden) were placed in the tibia of 15 adult rabbits, each with three to four exposed threads per implant. A PTFE membrane was placed over the test fixtures, covering the threads and 5 to 8 mm of the adjacent bone. The muscle and periosteum were replaced, adapted, and sutured. The control fixtures were not covered with a membrane. After healing periods of 6, 9, and 15 weeks, the specimens were removed en bloc and evaluated grossly and histolog-

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ically. Newly formed bone was evaluated with respect to structure, quality, and quantity. The results showed that all exposed threads of the titanium implants were covered with newly formed bone at a uniform thickness, even as early as 6 weeks. New bone formation also was seen in the control areas, although to a much lesser extent than the test areas. It was shown that by placing an inert membrane with an appropriate pore size, which hindered the penetration of undesirable cells, a space was created that permitted the entrance of osteogenic and angiogenic cells from the adjacent bone marrow to populate the area and proliferate. It also was recognized that the amount of new bone formed was contingent upon the amount of space created by the membrane. Further studies by Dahlin et al. [11] continued to provide support for the principle of guided bone regeneration in the regeneration of bone for clinical application. Osseous defects surgically created in the maxilla and mandible of monkeys, which were covered by a barrier membrane, evidenced complete regeneration of bone. An experimental study by Dahlin et al. [10] investigated the use of guided bone regeneration in an osseous defect in the calvaria of rats, a lesser osteogenic site than studied previously. Rather than treat the defect with an alloplastic material, which is a more commonly used graft material for cranial defects, the researchers grafted the site with autogenous bone. Seventy-two rats were treated with six alternative procedures for the healing of 8-mm, fullthickness, calvarial defects. A PTFE membrane was used in external placement only or as a combined internal and external placement. A high degree of bone regeneration was seen in all treatment groups regardless of the healing period, except for the control defect, which exhibited only slight, reparative osteogenesis. The principle of guided bone regeneration to support osteogenesis was again realized. This study demonstrated its use in larger osseous defects. The autogenous bone chips also functioned to maintain the space created between the inner and outer membrane. Without the preservation of the space, the membranes would have collapsed and prevented the ingrowth of osteogenic cells from the adjacent bone marrow. Finally, of clinical significance, Linde et al. [25] showed that new bone could be created at anatomic sites where bone is normally not present using expanded polytetrafluorethylene (e-PTFE) membranes of varying stiffness. Domes, 5 and 8 mm in diameter, made of e-PTFE (Gore-Tex) of differing degrees of stiffness were surgically placed over an area of denuded calvarial bone in rats and covered by the scalp. Histologic results demonstrated neogenesis

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of bone; however, the degree depended on the membrane stiffness and porosity and the length of the healing time. This study clearly demonstrated yet another use for the osteopromotive membrane technique, and it critically demonstrated practical details of the technique, such as the adequate formation of a blood clot, membrane stiffness and porosity, and the length of healing periods.

The osteopromotion principle Osteopromotion is a term used to describe the technique of physically placing a barrier over an anatomic, osseous site, thus segregating the site from the surrounding soft tissues, which are known to interfere with osteogenesis or the formation of new bone. The use of this term is germane to the terminology that describes bone forming mechanisms, such as osteoinduction, osteoconduction, osteogenesis, and osseointegration. As discussed previously, the use of the barrier membrane enhances complete osteogenesis by preventing the rapid ingrowth of fibroblasts into a bony defect and promoting the migration of osteogenic cells from the adjacent bony edges or bone marrow into the defect in an unimpeded fashion (Fig. 1). Leaving the barrier membrane in place for an extended period of time and sealing off the bony defect permits uninterrupted osteogenesis to occur and allows maturation of the newly formed bone. Of interest is how this mechanism works. Urist and McLean [48] stated that an inverse relationship exists between fibrinogenesis and osteogenesis. Formation of new bone depends on the proliferation of the endosteal and periosteal cells with osteogenic potential to form bone within defect without the deleterious effects of the competing and rapidly proliferating fibroblasts. It is also known that fibroblasts may inhibit osteoblast differentation by contact inhibition and by the release of soluble factors, such as prostaglandins. In vitro experimentation using cocultures by Ogiso et al. [35] has shown that rat skin fibroblasts and human periodontal ligament fibroblasts inhibited the formation of mineralized bone nodules in rat bone marrow stromal cell cultures. When bone marrow cells were cultured in a media from fibroblast cultures, inhibition of osteoblast function was observed, presumably through the production of inhibitory factors such as prostaglandins by the fibroblasts. Finally, growth factors and cytokines, which are potent mitogens for marrow stromal cells, are produced by these cells and stored in the extracellular matrix of bone. These factors act on osteoprogenitor cells in several ways to promote cell

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Fig. 1. Osteopromotion principle. The placement of a barrier membrane to isolate a bone defect from the surrounding soft tissue, permitting the exclusive migration of osteogenic cells from the adjacent bone. (Courtesy of W.L. Gore & Associates, Inc., Flagstaff, AZ, with permission.)

differentiation which results in osteogenesis, or they act as inhibitors of osteogenic differentiation. The role of the membrane and the success of the technique may be not only to act as a physical barrier preventing the invasion of fibroblasts into the defect but also to eliminate the effect of contact inhibition by the fibroblasts and to concentrate the previously mentioned growth factors within the area of new bone formation. The membrane itself may be osteoconductive. Circumstantial evidence exists in several studies that demonstrate the development of a thin layer of trabecular bone directly adjacent to the membrane [26].

Promotion of bone regeneration A brief discussion of the biology of bone regeneration is appropriate to understand further the principle of osteopromotion. The repair of a bone defect, as opposed to a fracture, is a preferred model to study bone regeneration. A surgically created bone defect offers a more controlled situation for experimentation because it is not subject to the variances of mechanical instability or an erratic blood supply [41]. A major clinical problem is healing of a critical size defect, which is known as the smallest size of an intraosseous wound that would not heal spontaneously during the lifetime of the individual [43]. Experimental research continues to evaluate the healing parameters of such a defect with the goal of potentially eliminating the need for bone grafting procedures altogether. Schenk [40] studied the pattern of bone regeneration in membrane-protected, critical size defects in the mandible of four adult male foxhounds. Extrac-

tion of three mandibular premolars bilaterally created a 50-mm edentulous space. Four months after the extractions, membrane surgery was performed. Two through-and-through rectangular defects were created bilaterally that measured approximately 8 mm vertically, 12 mm mesiodistally, and 10 mm buccolingually at the most inferior aspect of the defect. Each dog received (1) one standard, e-PTFE membrane (Gore-Tex Augmentation Material), (2) two prototype reinforced e-PTFE membranes (r-GTAM) that had been preformed into an arch shape, and (3) one nonmembrane control site. The membranes were sized to cover the defect and fit beyond the bony margin by 2 to 3 mm. The membranes also were secured to the alveolus with fixation screws. Of importance is the fact that intravenous blood was injected into all sites inferior to the membrane or inferior to the periosteum in the control site to confirm the formation of a stable blood clot within the isolated space. Primary wound closure was carried out. Healing was permitted for 2 months in two dogs and 4 months in the remaining two dogs. The specimens were harvested, radiographed, and evaluated by histologic review and histomorphometric analysis. Each of the four control sites demonstrated incomplete bone regeneration restricted to the bony margins of the defect. Copious scar tissue formation also was present within the defect. A deep indentation persisted along the alveolar crest that was partially filled by the collapsed mucosa, a finding not observed in the membrane-covered defects. The microscopic pattern of bone formation occurred, as demonstrated in the 2-month and 4-month healing specimens, in three general categories or phases. The most immature bone was woven bone, a random orientation of col-

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lagen fibrils along the margins of the surgically created defect. Next, parallel-fibered bone was deposited on the surface of the primary spongiosa as reinforcement. Finally, lamellar bone was seen, which represented the most mature phase of bone regeneration. Bone remodeling also was observed in this phase. Significant conclusions are evidenced by this study. First, this study confirms conclusions of previous experiments that demonstrated bone regeneration in membrane-protected defects. Second, the environment created by the membrane permitted the regeneration of bone by physically supporting the overlying soft tissue and preventing collapse of the soft tissue into the defect. It also protected the blood clot from surrounding soft tissue invasion while maintaining a space into which osteogenic cells could migrate and possibly seclude local growth factors and bone-promoting substances. Third, bone regeneration occurred, as mentioned earlier, in three stages. Fourth, complete bone regeneration was not seen at 4 months which provoked questions regarding the required healing time for the regenerate and the physical properties of the membrane itself, particularly of the resorbable type. Finally, the reinforced membranes maintained their original structure throughout the study as compared to the e-PTFE membrane, which showed minor deformation from the overlying soft tissue [40]. Several questions developed as a result of this study: Could an implant achieve osseointegration when placed in bone regenerated in a membraneprotected defect? Once loaded, would the implant stimulate bone remodeling and maturation of the regenerated bone? What is the fate of membraneprotected, regenerated bone that does not receive functional forces applied from an implant? Buser et al. [8] performed an inquiry experimentally to evaluate those questions and published the results in 1995. Fifteen nonsubmerged implants were placed bilaterally in regenerated bone created in membrane-protected defects in the mandible of five adult foxhounds. A 6-month healing period occurred between the surgical creation of the defects and the placement of the implants. Clinical and radiographic evaluation confirmed functional ankylosis of all 15 implants 3 months after placement. Subsequently, 8 implants were restored and functionally loaded for 6 months, leaving 7 implants unrestored. At the termination of the study, 9 months after implant placement, the histologic analysis demonstrated direct bone to implant contact for all 15 implants. It may be concluded that bone regenerated under a barrier membrane responds to implant placement similarly to native, nonregenerated bone by stimulating bone mat-

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uration and remodeling. The newly formed bone also is capable of sustaining functional loading. Interestingly, the regenerated bone that served as a control (no implant placed) demonstrated bone atrophy beneath the membrane. The results of this study are clearly relevant and applicable to the reconstructive surgeon. Pilot studies in humans, case reports, and multicenter studies continue to support the use of guided bone regeneration using a barrier membrane. For example, Becker et al. [3] conducted a prospective multicenter clinical study that determined the predictability for implants placed into immediate extraction sockets and augmented with e-PTFE barrier membranes. A total of 49 implants were placed with e-PTFE membranes. Three implants were lost at abutment connection surgery. Initial and final defect measurements and thread exposure were compared. Patients were followed up to 1 year after implant loading. The average bone formation for membrane-retained sites was 4.8 mm, whereas the average bone formation for sites in which the membranes were prematurely removed (20) was 4 mm ( P < 0.0001). At stage 2 surgery, an average of 0.6 threads ( P < 0.001) were exposed for the membrane retained sites and 2.6 threads for the early removal sites. Forty-five pairs of nonstandard radiographs were evaluated for bone loss after implant loading (7.5 months). The average mesiodistal bone loss averaged 0.72 mm. The results of this study demonstrated that by securing an e-PTFE membrane over an endosseous implant placed into an immediate extraction socket, substantial amounts of bone formation occurred adjacent to the implant. Sites at which the membranes were retained until stage 2 surgery also had the greatest amounts of bone formation.

Membrane design criteria As seen in the many studies using barrier membranes of differing materials, an evolutionary process has occurred in an attempt to develop the ideal membrane that satisfies several design criteria. A biomaterial is defined as a nonviable material used in a medical device that is intended to interact with biologic systems [4]. A biomaterial used within the body to address a particular function must fulfill two requirements, safety and efficacy. Safety is evaluated by a series of in vitro and in vivo experiments that assess various aspects of biocompatibility. For the biomaterial to be effective, it must meet criteria based on the biology of the surrounding environment and effectively perform specific functions [47].

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Hardwick et al. [18,39] have proposed design criteria for guided bone regeneration that include several factors in addition to the obvious requirement of a passive, physical barrier. First, the membrane must be constructed of an accepted biocompatible material. This criterion infers that the interaction between the material and the tissue will not adversely affect the individual by an undesired immune response or adversely affect the proposed outcome. The material also must be able to interact with the bone and soft tissue to allow for adaptability and stabilization. Second, the membrane must possess the ability to exclude tissues or cells. The membrane must contain structural elements that render the membrane impervious to the overlying soft tissue and associated fibroblasts. Naturally, the membrane also must be permeable to tissue fluids and nutritional macromolecules. Resistance to bacterial contamination is also important and necessary. Should the membrane become exposed to the oral environment, it must have the ability to protect the underlying bone regenerate from being contaminated and destroyed because of an inflammatory process. Third, another necessary criterion of the membrane is its ability to accommodate a space into which the regenerate can develop. This concept reflects the geometry and the volume of the space to be regenerated with bone. The membrane structurally must contain elements that provide stiffness to the membrane which allows the membrane to maintain its determined shape and resist collapse from the pressure of the overlying soft tissues or masticatory forces. This space-making characteristic also must exist throughout the entire healing period. An inert, nondegradable membrane processes this characteristic as opposed to membranes constructed from degradable material (Fig. 2).

Fourth, the membrane must permit tissue integration, a process that allows or even encourages the ingrowth or surface bonding of the overlying tissues. This feature promotes stabilization of the wound and inhibits epithelial migration along the tooth/gingival flap junction or implant/gingival flap junction. At the same time, the membrane must maintain its structural integrity to provide for cell exclusion. Finally, a guided tissue regeneration membrane must be a device that is easy to manipulate clinically and manage postoperatively. Easy use of the membrane is important when one considers that the membrane requires physical sizing and trimming to adapt it to a specific space and possibly is secured into its final position with sutures or fixation screws. If, however, it is necessary to remove the membrane because of untoward postoperative complications or as required in the second-stage surgery, the integrity of the membrane must be such that it may be removed without disintegrating [18,39]. As the general design criteria have developed and matured for guided bone regeneration in conjunction with endosseous implants and for the regeneration of large osseous defects, manufacturers have introduced many diverse barriers into the professional arena. A major distinction made between these membranes is their biologic behavior: the prototypical nonresorbable membrane and the newer absorbable, or bioresorbable, membrane.

Nonresorbable barriers General considerations Nonresorbable barriers were the first devices approved for clinical use. Because they maintain their structural integrity, their inherent desirable qualities

Fig. 2. Spacemaking (A) and nonspacemaking (B) principle. Spacemaking as a desirable quality of barrier membranes, as opposed to nonspacemaking membranes. (Courtesy of W.L. Gore & Associates, Inc., Flagstaff, AZ, with permission.)

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are exhibited throughout their existence in function. Essentially, this provides the operator with complete control as to the time frame of use and minimizes unpredictable events. Unfortunately, this type of barrier membrane must be removed which necessitates a second surgical procedure. Patient acceptance, cost, timing, and possible morbidity of the second surgery become considerations when choosing this type of membrane. Material considerations Most nonresorbable devices are made from PTFE or e-PTFE. PTFE is a fluorocarbon polymer that is inert and biocompatible. It is nonporous, prevents tissue ingrowth, and does not elicit a foreign body reaction in humans [49]. Expanded PTFE is chemically identical to PTFE, is also inert, and, when constructed to do so, allows for the interaction of desirable soft tissue for stabilization. Permeability is present to allow nutrients and fluids to filter through the membrane. It has a porous microstructure of solid nodes and fibrils. Membrane characteristics such as porosity, thickness, and rigidity may be adjusted to satisfy a specific need. This material has been used in vascular surgery and experimentally for many years [15,19]. The Gore-Tex e-PTFE barrier membrane features two structural designs that satisfy the desired criteria

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mentioned previously. An open microstructure collar that is only 1 mm thick, the membrane is of low density and is 90% porous (100 – 300 mm between nodes). This structure allows for connective tissue ingrowth and permits stability yet inhibits epithelial migration. The remainder of the membrane is partially occlusive, is stable, and maintains the space needed for bone regeneration. It is structurally 0.15 mm thick, is of higher density, and is 30% porous (less than 8mm between nodes) (Figs. 3 and 4). Observations from histologic specimens and clinical findings in humans continue to support the use of this membrane for the successful regeneration of bone for osseous defects, in the treatment of periodontally diseased teeth, and for staged and immediate implant placement in a volume deficient, osseous site [7,23,42]. The use of e-PTFE barrier membrane has been associated with minimal postoperative complications such as pain, purulence, swelling, soft tissue sluffing, and premature membrane exposure, which, if unmanageable, necessitates the early removal of the membrane. It has been shown that continued exposure of the barrier membrane requires surgical removal to prevent bacterial contamination of the underlying bone regenerate [46]. The Gore-Tex e-PTFE membrane also has been modified with titanium reinforcing struts placed between the two layers of expanded PTFE. This modification produces a membrane with identical surface

Fig. 3. e-PTFE membranes. Different configuration of e-PTFE membranes designed for guided bone regeneration. Note the inner (i) and outer (O) portions of the membrane. (Courtesy of W.L. Gore & Associates, Inc., Flagstaff, AZ, with permission.)

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Fig. 4. e-PTFE membrane is placed over bone grafting material for lateral ridge augmentation.

properties but with greater rigidity, which improves the space-making characteristic of the membrane. As discussed previously, Hurley, Melcher, Boyne and Bassett [2,5,20,29] fashioned sheets of celluloseacetate to function as a barrier membrane in experimental studies. Millipore membrane filters are a biologically inert mixture of cellulose-acetate and cellulose nitrate with varying pore sizes that range from 0.025 to 8 mm. Nyman et al. [33,34] also used the filter paper to study guided tissue regeneration. The device persisted for 6 months after implantation. This material is also described as a tested, nonresorbable membrane.

Bioresorbable barriers General considerations The use of bioresorbable membrane for guided bone regeneration precludes the necessity for a second surgical procedure, a fact attractive to many patients and surgeons. By their very nature, resorbable barriers offer limited control over the length of

application. Their functional requirements in guided bone regeneration are equivalent to nonresorbable membranes, however. A significant trend is seen in the switch from nonresorbable to resorbable membranes in guided bone and tissue regeneration, which has led to increasing research focused on creating the ideal resorbable membrane with the desired design criteria mentioned previously. The issues of space maintenance as determined by membrane rigidity, degradation rate, and degradation byproducts continue to be a source of inquiry for researchers. Because of their inherent nature, these membranes elicit tissue reactions that may influence wound healing and their overall effectiveness. Ideally, these reactions should not compromise the desired regenerative outcome or the general well-being of the host. The end products of degradation are cited as the main factors that influence the safety and effectiveness of the membrane. Absorbable materials used for guided bone regeneration are divided into two broad categories: natural products and synthetic materials. A brief overview of the definitions of bioabsorbable, biodegradable, and bioresorbable is pertinent to this article to prevent misconceptions regarding

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resorbable membrane activity. These definitions were discussed by Vert [50]. Biodegradable refers to solid polymeric materials and devices that break down as a result of macromolecular degradation with dispersion in vivo, but there is no proof of elimination from the body. Fragmentation or other degradation of byproducts occurs that may move them away from their site of implantation but not from the body. In contrast, bioresorbable refers to a solid polymeric material that can degrade and further resorb in vivo, and is eliminated through natural pathways of filtration or by being metabolized. This process reflects the fact that the material is totally eliminated without residual side effects. Bioabsorbable is the dissolution of solid polymeric materials in body fluids without a polymer chain cleavage or molecular mass decrease. A bioabsorbable polymer may be bioresorbed if the dispersed macromolecules are excreted. Material considerations Natural products Several medical devices have been manufactured from collagen because of its biologic and physical properties. Easily available, collagen represents approximately one third of total body proteins. The 19 types of collagen, of which type I is the most common, differ in their structure and macromolecular arrangement [36]. Collagen is easily the most studied natural polymer. In general, it is derived from bovine intestine, hence the name gut. It is also derived from bovine tendon, skin, and sheep intestine. Isolation and purification to minimize an-

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tigenicity follow one of two ways: enzymatic preparation and chemical extraction. The collagen may be processed into several manufactured devices, such as gels, hemostatic sponges, filaments, membranes, and sutures, for medical use. A common processing technique is cross-linking with (glutar)aldehyde to decrease toxicity of the final product and change the degradation time and increase the tensile strength of the collagen fibers. Gut or other collagen materials are degraded by sequential attack by lysosomal enzymes and collagenase from infiltrating macrophages and polymorphonuclear leukocytes and continue on a course of native enzymatic pathways, which leads to peptide fragments and amino acid residues [21,47]. Although concerns have been raised about induced anticollagen antibodies, particularly with the use of injectable collagen and the potential transmission of bovine spongiform encephalopathy from injected bovine products, a US Food and Drug Administration panel has determined these products safe for their approved function [35]. Bio-Gide membrane (Ed. Geistlich Sohne AG, Wolhusen, Switzerland) is a resorbable pure porcine collagen membrane manufactured as a bilayer structure used in guided bone and tissue regeneration. The membrane is made of pure type I and type III collagen. Maintenance of the barrier function is reported for a period of 24-weeks (Fig. 5). BioMend Extend (Integra LifeSciences Corp. for Sulzer Calcitek, Inc, Carlsbad, CA) is also a resorbable collagen membrane indicated for use in guided tissue and bone regeneration. It is fabricated from

Fig. 5. Bioabsorbable membrane is placed over bone grafting material for lateral ridge augmentation.

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bovine deep flexor (Achilles) tendon and is pure type I collagen. According to the manufacturer, BioMend is completely absorbed in 4 to 8 weeks. The use of bovine pericardium is reported in the literature for various medical procedures, including, but not limited to, guided bone regeneration [14,22]. Non – crossed-linked (lyophilized) and cross-linked membranes with glutaraldehyde are available for use. Little information exists in the literature currently to support its use for guided bone regeneration in the oral cavity. A calcium sulfate barrier, Capset (Lifecore Biomedical, Chaska, MN), may be used as a barrier membrane over bone grafting material in periodontal defects and extractions sites. Sterile, medical grade calcium sulfate is mixed with an accelerating solution to form a moldable paste. The paste is then spread over the graft material and exhibits a working time of approximately 3 minutes. Resorption occurs approximately 4 weeks after placement. Few reports exist in the literature concerning the effectiveness of this material as a barrier membrane. Other available natural products that have been tested for guided bone regeneration include dura mater and laminar bone [28,52]. Dura mater, an irregular network of collagen fibers, is obtained from cadavers, processed to remove antigenic and pyrogenic activity, and then lypholized and sterilized. A limited inflammatory response is evident histologically in human biopsies, and most of the material is resorbed within 6 weeks of surgery. Transmission of Creutzfeldt-Jakob disease is a risk, although an unlikely one, that must be considered when using cadaveric dura mater [51]. Laminar bone (Lambone, Pacific Coast Tissue Bank, Los Angeles, CA) a 20 to 100 mm or 100 to 300 mm thick strip of demineralized, freeze-dried bone, has been used as a guided tissue regenerative device. It is an allograft recovered from a carefully screened human donor. Limited information is available in the literature regarding its effectiveness as an osteopromotive device, potential osteoinductive and osteoconductive activity, resorption time, and ease of use as a barrier membrane [28,37]. Synthetic products Synthetic absorbable devices for medical use have been used for several decades and have undergone extensive testing in in vivo and in vitro studies [6,16]. Bioabsorbable polymers are a special class of polymers in which the material serves a function, is gradually broken down, is metabolized, and is then eliminated from the body. Bioabsorbable materials undergo a two-phase degradation process in the body.

Phase I, primarily physical in nature and occuring as a result of water molecules hydrolytically attacking the polymeric chemical bonds, decreases the strength of the material. Phase II is the physiologic response of the body. Macrophages phagocytize the fragments and metabolize them to common substances such as water and carbon dioxide via the citric acid cycle. At this stage, the mass of the material quickly disappears. The degradation rate depends on the pH, mechanical strain, enzymes, and bacteria. Material composition also varies the rate. The materials most commonly used are poly (a-hydroxyacids), that include poly(lactic acid), poly(glycolic acid), and their copolymer(s) poly(glycolide-lactide). Poly(glycolic acid) and poly(lactic acid) are manufactured by polymerization of the monomers and are widely used for sutures and drug controlledrelease device. As mentioned previously, varying the composition or the percentage of poly(L-lactide) to poly(glycolic acid) modifies their degradation rates. Poly(glycolic acid) degrades more quickly of the two, and poly(L-lactide) is the more stable in vitro [47]. Control of the degradation process as it relates to the function of the material in guided tissue regeneration is a continued goal in the manufacturing process of the ‘‘perfect membrane,’’ one that exhibits appropriate strength, achieves biomechanical demands, degrades in a predictable fashion, and does not interfere with bone regeneration and healing. One disadvantage associated with homopolymers and copolymers has been a rather ill-defined inflammatory and foreign body reaction. Treatment has included incision and drainage of the site, surgical debridement, or ultimately, the complete removal of the material [31,44]. Polyglactin 910 mesh (Vicryl, Ethicon, Inc, Sommerville, NJ) is an inert, nonpyrogenic, nonantigenic material prepared from a synthetic absorbable copolymer of glycolide and lactide derived respectively from glycolic and lactic acids. It is known to elicit a mild tissue reaction during absorption. Polyglactin 910 has been used as a membrane experimentally for guided tissue regeneration in a study in rabbit tibia in a non-critical size defect model [1]. According to the manufacturer, subcutaneous implantation studies in rats indicate that the absorption of the mesh material is minimal until approximately 6 weeks after implantation and is complete in 60 to 90 days. Gore Resolute XT (W.L. Gore & Associates) regenerative material is a 100% synthetic, bioabsorbable membrane. The membrane is composed of the synthetic polymers of polyglycolic acid, polylactic acid, and trimethylene carbonate. It is reported to offer 8 to 10 weeks of barrier function.

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Gore Osseoquest (W.L. Gore & Associates) regenerative material is a 100% synthetic bioabsorbable membrane that maintains a barrier function for 6 months and is substantially resorbed in 12 to 14 months. It is composed of polyglycolic acid, polylactic acid, and trimethylene carbonate. It is indicated to function as a barrier membrane for the regeneration of bone in a large osseous defect. Several investigative studies have been reported in the literature within the past 8 years that evaluate bioabsorbable membranes in guided tissue regeneration and compare these membranes to the effectiveness of a nonabsorbable membrane. Zellin et al. [53] published an article in 1995 comparing the osteopromotive potential of ten different biodegradable and nonbiodegradable membrane materials. Critical size defects were created bilaterally in the mandible of adult rats and randomly covered with different types of membranes. After 6 weeks of healing, an evaluation was performed using light microscopy. The authors concluded that different membranes differ significantly in their osteopromotive efficacy, even if closely related chemically. The study also showed that membranes developed for guided tissue regeneration in periodontal defects may not be adequate to promote osteogenesis [11]. A continuing concern with the use of a bioabsorbable membrane is its ability to physically retain its space-making characteristic for an adequate period of time while new bone is forming. Lundgren et al. [27] reported that it is advisable to use a supporting material to prevent barrier collapse because bioabsorbable barriers may not possess enough spacemaking properties for bone regeneration to occur in total. Exploring that premise, Mellonig, et al. [30] studied the efficacy of a bioabsorbable membrane in combination with decalcified freeze-dried bone allograft (DFDBA) for the correction of a dehiscence-type defect around endosseous titanium threaded implants placed in the mandibles of dogs. Six non-space-making defects created in the mandible were subsequently covered with a bioabsorbable membrane alone (Resolute), the same bioabsorbable membrane with DFDBA, an e-PTFE membrane (GTAM, WL Gore & Associates) combined with DFDBA, and finally, treatment alone with the replaced mucoperiosteal flap. Six months after surgery, each specimen was subject to histomorphometric analysis. The greatest percentage of bone contact (83.3%) with the implant was achieved by the GTAM and DFDBA, followed by the Resolute and DFDBA group (68.9%). Sites treated with Resolute alone or covered with the soft tissue flap alone demonstrated highly variable and less than desirable results. These

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results indicated that bone graft augmentation is necessary to achieve satisfactory results with bioabsorbable barrier membranes in the treatment of osseous defect. Such is not the case for a nonbioabsorbable membrane [30]. The greater efficacy of the e-PTFE nonresorbable membrane in promoting the formation of new bone in the treatment of defects around titanium dental implants placed in postextraction sites in humans was demonstrated by Simion et al. [45] when compared to the treatment of the same site by resorbable membranes of poly(lactic acid) and poly (glycolic acid). A recent investigation by Schliephake et al. [42] continued to show superior bone regeneration, particularly in height, under the e-PTFE membrane, as compared to the bone regenerated in peri-implant defects underneath a resorbable membrane. They also found that bone under the resorbable membrane showed sign of superficial resorption.

Clinical application From a clinical standpoint, the use of a bioabsorbable membrane for guided bone regeneration may be performed successfully by implementing the following technical guidelines. To prevent membrane collapse into the defect, the use of a filler material such as a bone substitute or a particulate autogenous bone graft is recommended. A synergistic relationship exists between the osteoconductivity of the graft and the membrane. Small perforations created in the cortical plate stimulate bleeding and access the marrow cavity, which contains cells with osteogenic potential. Finally, closure of the soft tissue flap in a tension-free manner is imperative.

Future goals Further research is necessary to develop a resorbable membrane that is distinctly biocompatible while exhibiting greater efficacy, specifically membrane longevity, to allow complete osteogenesis to occur. The relationship between membrane porosity, resorption kinetics, and mechanical properties must be defined further to create membrane equal in efficacy to the nonresorbable membrane for the treatment of osseous defects. Modifications to the membrane surface of inner structure must also be explored. The incorporation of an antibacterial substance is one example of a modification to evaluate or, in another approach, the addition of cytokines to

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enhance the rate of osteogenesis in a membraneprotected defect [9,38].

Summary As demonstrated in this article, the concept and clinical application of guided tissue and bone regeneration have been investigated exhaustively in the animal model and in human studies. Evidence clearly exists to support the successful clinical use of this technique for promotion of new bone formation in an osseous defect, although its mechanism remains somewhat undefined. Continued research into the techniques themselves, further applications, use of modifying factors, and the quest for a superior membrane are anticipated.

References [1] Aaboe M, Pinholt EM, Hjørting-Hansen E. Healing of experimentally created defects: a review. Br J Oral Maxillofac Surg 1995;33:312 – 8. [2] Bassett CA, Creighton DK, Stinchfield FE. Contributions of endosteum, cortex and soft tissue to osteogenesis. Surg Gynecol Obstet 1961;112:145 – 52. [3] Becker W, Dahlin C, Becker B, Lekholm U, van Steenberghe D, Higuchi K, et al. The use of e-PTFE barrier membranes for bone promotion around titanium implants placed into extraction sockets: a prospective multicenter study. Int J Oral Maxillofac Implants 1994;9: 31 – 40. [4] Black J. Biological performance of materials: fundamentals of biocompability. 2nd edition. New York: Marcel Dekker; 1992. [5] Boyne PJ, Mikels TE. Restoration of alveolar of ridges by intramandibular transposition osseous grafting. J Oral Maxillofac Surg 1968;26:569 – 76. [6] Brady JM, Cutright DE, Miller RA, Battistone G. Resorption rate, route of elimination and ultrastructure of the implant site of polylactic acid in the abdominal wall of the rat. J Biomed Mater Res 1973;7:155 – 66. [7] Buser D, Bra¨ger U, Lang NP, Nyman S. Regeneration and enlargement of jaw bone using guided tissue regeneration. Clin Oral Impl Res 1990;1:22 – 32. [8] Buser D, Ruskin J, Higginbottom F, Hardwick R, Dahlin C, Schenk RK. Osseointegration of titanium implants in bone regenerated in membrane-protected defects: a histologic study in the canine mandible. Int J Oral Maxillofac Implants 1995;10:666 – 81. [9] Chang C, Yamada S. Evaluation of the regenerative effects of a 25% doxycycline-loaded biodegradable membrane for guided tissue regeneration. J Periodontol 2000;71:1086 – 93. [10] Dahlin C, Alberius P, Linde A. Osteopromotion for cranioplasty. J Neurosurg 1991;74:487 – 91.

[11] Dahlin C, Gotttlow J, Linde A, Nyman S. Healing of maxillary and mandibular defects using a membrane technique: an experimental study in monkeys. Scand J Plast Reconstr Surg Hand Surg 1990;24:13 – 9. [12] Dahlin C, Linde A, Gottlow J, Nyman S. Healing of bone defects by guided tissue regeneration. Journal of Plastic and Reconstructive Surgery 1998;81:672 – 6. [13] Dahlin C, Sennerby L, Lekholm U, Linde A, Nyman S. Generation of new bone around titanium implants using a membrane technique: an experimental study in rabbits. Int J Oral Maxillofac Implants 1989;4:19 – 25. [14] Del Campo C, Konok G. Use of a pericardial xenograft patch in repair of resected retrohepatic vana cava. Can J Surg 1994;37:1:59 – 61. [15] Florian A, Colin LH, Dammin GJ, Collins JJ Jr. Small vessel replacement with Gore-Tex (expanded polytrafluoroethylene). Arch Surg 1976;111:267 – 70. [16] Gogolewski S, Pennings AJ. Resorbable materials of poly(L-lactide) 3. Porous materials for medical applications. Colloid Science 1983;261:477 – 84. [17] Gottlow J, Nyman S, Lindhe J, Karring T, Wennstro¨m J. New attachment formation in the human periodontium by guided tissue regeneration. J Clin Periodontol 1986;13:606 – 16. [18] Hardwick R, Scantlebury T, Sanchez R, Whitley N, Ambruster J. Membrane design criteria for guided bone regeneration of the alveolar ridge. In: Buser D, Dahlin C, Schenk RK, editors. Guided bone regeneration in implant dentistry. Chicago: Quintessence; 1994. p. 101 – 36. [19] Haydock D, Flint M, Hyde K, Reilly H, Poole A, Hill G. The efficacy of subcutaneous gortex implants in monitoring wound healing response in experimental protein deficiency. Connect Tissue Res 1988;17: 159 – 69. [20] Hurley LA, Stinchfield FE, Bassett CA. The role of soft tissue in osteogenesis: An experimental study in canine spine fusions. J Bone Joint Surg Am 1959;41: 1243 – 54. [21] Hutmacher D, Markus H, Schliephake H. A review of material properties of biodegradable and bioresorbable polymers and devices for GTR and GBR applications. Int J Oral Maxillofac Implants 1996;11:667 – 78. [22] James NL, Poole-Warren LA, Schindholm K. Comparative evaluation of treated bovine pericardium as a xenograft for hernia repair. Biomaterials 1991;12: 801 – 9. [23] Lang NP, Bra¨gger U, Hammermerle CHF. Immediate transmucosal implants using the principle of guided bone regeneration. Clin Oral Impl Res 1994;5:154 – 63. [24] Lekholm U, Zarb GA. Patient selection and preparation. In: Bra¨nemark P-I, Zarb GA, Albrektsson T, editors. Osseointegration in clinical dentistry. Chicago: Quintessence; 1985. p. 199 – 204. [25] Linde A, Alberius P, Dahlin C, Bjurstam K, Sundin Y. Osteopromotion: a soft-tissue exclusion principle using a membrane for bone healing and bone neogenesis. J Periodontol 1993;64:1116 – 28. [26] Linde A, Thoren C, Dahlin C, Sanberg E. Creation of

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In situ forming biomaterials$ Marsha Ritter Jones, MS, Phillip B. Messersmith, PhD* Biomedical Engineering Department, Northwestern University, 2145 Sheridan Road E310, Evanston, IL 60208, USA

In the last few decades, numerous advances have been made toward the development of in situ forming biomaterials. Research efforts in this area are fueled by the need for alternative and less invasive methods of administering biomaterials (i.e., functional materials that can be administered in a clinically acceptable manner, rapidly, and less invasively). The potential clinical applications of in situ forming biomaterials encompass many types of tissue repair, reconstruction, and drug delivery. The ideal in situ forming biomaterial can be manipulated easily and administered via syringe and needle into a tissue or organ. These materials are less invasive than conventional surgically implanted biomaterials and have the additional advantage of being capable of flowing into a body cavity or defect before solidification, which results in enhanced integration with native tissue. In situ forming biomaterials are typically formulated as a liquid or low viscosity gel and harden in vivo by various mechanisms. Although inorganic solids such as calcium phosphates may be formed by in situ processes, most in situ forming biomaterials are hydrogels. Most of these materials are suitable as drug delivery vehicles and for soft tissue repair, and many of the hydrogels used as in situ forming biomaterials can be used with cells and serve as scaffolding for tissue regeneration. This article covers existing and emerging approaches to in situ forming biomaterials. Although strictly speaking, most conventional dental restorative materials such as composites and adhesives and

$ This research was supported in part by NIH grants DE13030, DE12599, and DE13076. * Corresponding author. E-mail address: [email protected] (P.B. Messersmith).

orthopedic bone cements also can be classified as in situ forming biomaterials, this article does not cover these materials because they have been extensively reviewed in the past. Instead, the article focuses on hydrogels and similar soft materials that are being developed for use as injectable materials and the potential clinical applications of these materials. The authors have elected arbitrarily to organize the review according to four broadly defined mechanisms used for triggering in situ formation of biomaterials: physiologic stimuli (e.g., temperature and pH), physical changes in the biomaterial (e.g., solvent exchange and swelling), chemical reactions (e.g., enzymatic, chemical, and photoinitiated polymerizations), and self-assembly.

In situ formation based on physiologic stimuli Thermally triggered systems The use of biomaterials whose transition from sol to gel is triggered by an increase in temperature is an attractive way to approach in situ formations. The ideal critical temperature range for such systems is between ambient and physiologic temperature, such that clinical manipulation is facilitated and no external source of heat, other than that of the body, is required to trigger gelation. A useful system also should be tailorable to account for small differences in local temperature, such as might be encountered in an appendage, at the surface of the skin, or in the oral cavity. Three main strategies exist for engineering a thermally responsive sol-gel polymer system. Polymers with lower critical solution temperature (LCST) transitions between ambient and physiologic temperature have been used for this purpose. Such polymers undergo a hydrophilic-to-hydrophobic thermody-

1042-3699/02/$ – see front matter D 2002, Elsevier Science (USA). All rights reserved. PII: S 1 0 4 2 - 3 6 9 9 ( 0 2 ) 0 0 0 1 5 - 8

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namic transition upon heating through the LCST, which results in a sol-gel transition of an aqueous solution of the polymer. Self-assembly induced by thermal disorder-order transitions of block copolymer solutions can be exploited to drive sol-gel transitions. Finally, thermally responsive liposomes have been exploited recently to drive sol-gel transitions of several biomaterials. One of the most extensively investigated polymers that exhibits a useful LCST transition is poly(Nisopropyl acrylamide) (PNIPAAm) [42]. PNIPAAm is a water-soluble polymer below its LCST but is hydrophobic above the LCST, which results in precipitation of PNIPAAm from solution at the LCST. Copolymerization of NIPAAm with other monomers provides a route to tailoring LCST and other polymer properties [42]. For example, copolymerization with acrylic acid (AAc) reduces the LCST in a concentration dependent manner, whereas copolymerization with a bifunctional cross-linker such as methylenebisacrylamide (MBA) results in the formation of thermally responsive PNIPAAm gels that undergo reversible contraction and expansion when the LCST is traversed. Vernon et al. [52] have developed an injectable form of PNIPAAm that consists of high molecular weight linear polymer that gels with little syneresis when heated above the LCST. The material was suggested to be useful for drug delivery and cell encapsulation applications. Stile et al. [46] recently developed a weakly cross-linked form of PNIPAAm by copolymerizing NIPAAm with AAc and small amounts of MBA. The resulting weakly cross-linked hydrogel was injectable and became rigid at the LCST but exhibited virtually no volumetric shrinkage. They went on to study the use of P(NIPAAm-co-AAc) hydrogels, including peptide-modified hydrogels, as injectable scaffolds for tissue engineering [47]. Rat calvarial osteoblasts seeded into the peptide-modified hydrogels were viable for at least 21 days of in vitro culture, and cells were observed to spread more and demonstrated significantly greater proliferation when cultured within the peptide-modified hydrogels, as compared to control hydrogels. These peptide-modified P(NIPAAm-co-AAc) hydrogels serve as useful tools for studying cell-material interactions within three-dimensional structures and have the potential to be used as injectable scaffolds for tissue engineering applications. Pluronics are poly(ethylene oxide)-poly(propylene oxide)-poly(ethylene oxide) (PEO-PPO-PEO) triblock copolymers that are water soluble at low temperatures [2]. At high polymer concentration in water, the block copolymer molecules self-assemble

into micelles with a hydrophobic PPO core and a corona composed of hydrophilic PEO segments. The micelle structure of the block copolymers can be exploited for drug delivery by loading a hydrophobic drug into the micelle core; likewise, hydrophilic drugs can be loaded into the shell of the micelle. At concentrations above 15% to 20% in water, pluronics are fluid at low temperature but form a thermoreversible gel when heated as a consequence of a disorder-order transition in micelle packing [2], which makes these polymers suitable for in situ gelation. Esposito et al. [13] compared the efficacy of drug delivery using pluronic and myverol (glycerol monooleate) gels for periodontal disease treatment [8]. In vitro studies demonstrated that the pluronic gel released tetracycline more rapidly than myverol, which indicates that pluronic may be suitable for short-term release of drugs in the oral cavity. When tested in a clinical trial of 24 patients with periodontitis, the pluronic carrier performed better than myverol, reducing probing depth from 6.2 to 3.9 mm and reducing bleeding on probing from 75% to 17%. Bioadhesive polymers (e.g., carbopol, polycarbophil, and sodium alginate) have been added to pluronics to improve bioadhesivity, an important property when the materials are in contact with mucosal surfaces. In vitro studies have demonstrated that addition of bioadhesive polymers increased the strength and mucoadhesive properties of pluronic, with alginate having the greatest effect [40]. In in vivo studies using Sprague-Dawley rats, the mean residence time of propanolol, a beta-blocker, was greater for alginatepluronic suppositories than for the pure pluronic suppositories. The drug residence time was similar to that achieved by intravenous administration, which was believed to be caused by lack of hepatic first pass effects for the orally administered drug. Another triblock copolymer, poly(ethylene oxide)poly(l-lactic-co-glycolic acid)-poly(ethylene oxide) (PEO-PLGA-PEO), is a sol at room temperature but becomes a gel at body temperature [26]. Jeong et al. [26] examined the diffusion of hydrophobic and hydrophilic drugs from the gel and found that hydrophilic drugs diffused completely out of the gel before the gel dissolved. This was in contrast to hydrophobic drugs, which diffused much slower and were released primarily by dissolution of the gel. The PEO-PLGAPEO block copolymer has the added benefit of being resistant to cell adhesion because of the PEO segments and being biodegradable because of PLGA segment [24]. This polymer persisted in situ for approximately a month, which makes it ideal for long-term drug delivery [25].

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Fig. 1. Thermally triggered hydrogel formation from a suspension containing Ca-loaded liposomes (open circles), drug-filled liposomes (shaded circles), and alginate (lines). The Ca-loaded vesicles are designed to rapidly release Ca2 + in response to a change in temperature, crosslinking the aliginate to form a hydrogel containing entrapped drug-filled liposomes. (From the American Chemical Society.) [11]

An emerging strategy for thermally inducing in situ formation of biomaterials exploits liposomes to compartmentalize chemical triggers of gelation [8,31,55]. Cui and Messersmith [8] used calciumcontaining liposomes designed to release their contents when heated to 37
C to develop an injectable alginate fluid that gelled in less than 1 minute at 37
C (Fig. 1). Drug-filled liposomes added to the fluid before gelation were entrapped within the alginate gel network and released their contents at a rate that depended on lipid composition. In vitro studies of metronidazole released from in situ formed liposome/ alginate hydrogels suggested that this method could

be used for controlled release of drugs from the gel to the sorrounding environment (Fig. 2). In an extension of this concept, Westhaus and Messersmith [55] have demonstrated the use of liposomes for thermally triggered enzymatic gelation of protein-based hydrogels. In this case, the liposomes were loaded with calcium and combined with fibrinogen and human recombinant factor XIII, which is a calcium-dependent transglutaminase enzyme. When heated to 37
C, calcium released from the liposomes triggered activation of enzymemediated cross-linking of the protein into a hydrogel. Formation of fibrinogen hydrogels was achieved

Fig. 2. Release of metronidazole from thermally gelled liposome/alginate hydrogels. Sealed dialysis bags containing one part by volume Ca-loaded liposomes, one part drug-loaded liposomes (DPPC or 15% DMPC), and three parts Na-alginate (2%) were inserted into 37
C buffer at time zero. The control was a metronidazole-infused Ca-alginate hydrogel (2%). Triangle = caalginate, square = 15% DMPC, circle = DPPC. (From the American Chemical Society.) [11]

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in less than 10 minutes using this strategy, and the gels that form are expected to have a microstructure different from fibrin-based gels formed by clotting of fibrinogen by thrombin. Finally, Murphy and Messersmith [31] have reported on the use of thermally responsive liposomes to form calcium phosphate minerals for remineralization of dentin and enamel. In this case, liposomes were loaded separately with calcium and phosphate and mixed together to form a vesicle suspension highly supersaturated with respect to calcium phospate minerals. When the liposome suspension was heated to 37
C, liposomes released their contents, and calcium phosphate mineral was formed. In vitro experiments in which thermally triggered mineralization was performed in the presence of dentin and enamel substrates demonstrated the deposition of mineral onto the tissue surfaces. This strategy could be useful in mouthrinse form for rapid remineralization of incipient dental caries at oral temperatures. pH triggered systems Another mechanism of in situ formation based on physiologic stimuli is the formation of gels induced by pH changes. Although the pH mechanism has not been exploited widely for delivery of materials to the oral cavity, there is great potential for designing materials that gel or modulate their properties in response to oral pH changes, and several examples of pH-sensitive materials exist that have been applied to other anatomic locations. For example, in situ hydrogel formation is particularly useful in ocular drug delivery. Drugs formulated as liquid solutions have several limitations, including limited bioavailability and propensity to be rapidly removed by tear fluid [27]. Kumar and Himmelstein [28] sought to minimize these factors and maximize drug delivery by making a poly(acrylic acid) (PAA) solution that would gel at pH 7.4. The authors found that at concentrations high enough to cause gelation, however, the low pH of a PAA solution would cause damage to the surface of the eye before being neutralized by the lacrimal fluid. This problem was solved partially by combining PAA with hydroxypropyl methylcellulose (HPMC), a viscous enhancing polymer, which resulted in a pH-responsive polymer mixture that was a sol at pH 4 and a gel at pH 7.4. Mixtures of poly(methacrylic acid) (PMA) and poly(ethylene glycol) (PEG) also have been used as a pH sensitive system to achieve gelation [17]. This mixture is soluble in a water-ethanol mixture below pH 5.7. Upon infusion into the body, the ethanol diffuses into the peripheral tissue and the solution

gels. The gel is believed to result from entanglement of polymer strands, van der Waals forces, and hydrogen bonding between the PMA and PEG. The study found that a PEG (MW 18.5 k) and PMA (MW 15 k) in a ratio of 1:2 was the best mixture with the best gelling performance. Slow dissolution of the polymer gel above pH 4.8 occurred as a result of dissociation of the acid groups on PMA and consequent disruption of the hydrogen bond network. The researchers were able to show that a gel could be maintained for a longer period of time by the addition of citric acid to the mixture before gelation, which caused the pH of the mixture to remain low for a longer period of time.

In situ formation based on physical mechanisms Swelling In situ formation also may occur when a material absorbs water from the surrounding environment and expands to occupy a desired space. One such substance is myverol 18-99 (glycerol mono-oleate), which is a polar lipid that swells in water to form lyotropic liquid crystalline phase structure [13,15]. It has some bioadhesive properties and can be degraded in vivo by enzymatic action [13]. Geraghty et al. [15] found that the gel’s adhesive and cohesive properties depended on water composition. Although gels were adherent to poly(methyl methacrylate) and mucosal surfaces, the presence of water on the mucosal surface was found to hinder the formation of adhesive bonds. One proposed use for myverol 18-99 is as a drug delivery vehicle. Esposito et al. [13] used a tetracycline-containing mixture of 90:10 myverol:water for the treatment of periodontal disease. In a 4-week clinical trial, a myverol carrier was shown to reduce the probing depth and bleeding on probing, which is an indication of periodontal stability. Diffusion This method involves the diffusion of solvent from a polymer solution into the surrounding tissue and results in the precipitation or solidification of the polymer matrix. N-methyl pyrrolidone (NMP) has been shown to be a useful solvent for such systems [30]. To treat intrabony and furcation defects and regeneration of the periodontium, Rosen and Reynolds [38] studied the use of resorbable poly(DLlactide) in combination with a composite graft and tetracycline. The polymer was dissolved in NMP and used to form a gel in the periosteal space. The polymer

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served to cover the soft tissue and completely seal the defect from epithelial and gingival connective tissue migration. The intrabony lesions had a decreased probing depth of 60% after 6 months from pretreatment states, and the gel was completely absorbed. A poly(lactide-co-caprolactone) copolymer in NMP or dimethyl sulfoxide also has been studied as a drug delivery system [39]. Ismail et al. [22] used a NMP/ ethanol/water solvent mixture to compare the use of PMA:PEG to that of an HPMC:carbopol mixture for plasmid DNA delivery. The PMA:PEG polymer mixture was buffered at pH 7.2 and dissolved in a mixture of NMP, water, and ethanol. The release of plasmid DNA from the PMA-PEG gel was low when compared to the HPMC:carbopol mixture. As with the HPMC:carbopol mixture, there was an increased release of the plasmid DNA with decreasing pH.

In situ formation based on chemical reactions Chemical reactions that result in in situ gelation may involve precipitation of inorganic solids from supersaturated ionic solutions, enzymatic processes, chemical cross-linking reactions, and photoinitiated processes. Some examples of chemical cross-linking reactions that fall into these categories have been discussed previously and include transglutaminasemediated cross-linking of proteins and peptidemimetic polymers into hydrogels, ionic cross-linking mechanisms such as in the case of calcium – crosslinked alginate, and chemical or photochemical initiation of free radical polymerization of acrylatecontaining polymers. Ionic cross-linking Aqueous solutions of sodium alginate have been investigated widely for in situ formation. Alginate is a block copolymer of mannuronate and guluronate, and the formation of alginate hydrogel relies on the ionic cross-linking of guluronate blocks by multivalent cations. Although extensively used as a dental impression material that solidifies when mixed with soluble calcium salts, only recently have reports emerged of attempts to use injectable forms of alginate for the purposes of drug delivery, tissue repair, and tissue engineering [7,9,14,32]. Alginate is viewed as a good scaffold for maintaining the phenotype of chondrocytes in culture [19]. Injectable strategies are being developed for use of alginate in transplantation and engraftment of chondrocytes for plastic and reconstructive surgery; most rely on diffusion of calcium ions from the surrounding fluid

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into a viscous alginate solution, which results in formation of calcium alginate hydrogel. Several studies have shown that cartilage-like tissue can be formed when chondrocytes are injected into a tissue and entrapped by in situ formation of an alginate matrix [7,9,14,32]. Fragonas et al. [14] have shown that an in situ gelled alginate suspension of chondrocytes could be used to heal cartilage defects. In vivo studies using New Zealand white rabbits demonstrated the formation of cartilage-like tissue as shown in Fig. 3, which is a hematoxylin and eosin stained section of the control and the alginate implant at 1, 2, and 4 months, which depicts the cartilage regeneration in the alginate implants. Paige et al. [32] implanted chondrocytes in an injectable alginate carrier into nude mice and found histologic evidence of cartilage-like tissue formation within 6 weeks. Diduch et al. [9] implanted marrow stromal cells in alginate and other matrices (type 1 collagen and agarose) to investigate the repair of osteochondral defects. In vitro studies compared the ability of these cells to produce aggrecan and type II collagen and showed that the production of type II collagen and aggrecan was greatest in agarose followed by alginate. The cells in alginate maintained their shape and size better than the other two matrices, however. The size, shape, and consistency of the alginate gel were also better maintained over time compared to the collagen and agarose matrices. In vivo studies involved the implantation of marrow stromal cells in an alginate carrier into a cartilage defect in rabbits and showed a restoration of the organization of the subchondral bone within 2 months of implantation. Cohen et al. [7] have explored the use of injectable alginate as an ophthalmic drug delivery system. In vitro studies using stimulated tear fluid and calcium-containing solutions suggest that strong gels can be achieved with relatively low concentration (0.5% – 1.5%) of alginate. Strong gels are necessary for this application because of the need to resist shear stresses induced by eye motion and blinking. The authors also investigated the effect of alginate composition on the gelation properties. These experiments demonstrated that alginates with high guluronate content gelled more rapidly and formed stronger gels than alginates with low guluronate content. Release of pilocarpine from these gels occurred slowly over a period of 24 hours. In vivo studies were conducted with pilocarpine and alginate gels to investigate the effect of pilocarpine release on intraocular pressure of rabbits with glaucoma. Prolonged release of drug from the high guluronate content gel reduced intraocular pressure

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Fig. 3. Hematoxylin-eosin stain of in situ formed aliginate implants embedded with chondrocytes at 1 (A), 2 (C), and 4 (E) months. The corresponding controls which were allowed to heal naturally (B, D, and F) (Original magnification  25) (From Elsevier Science.) [14]

for up to 10 hours after administration, whereas the low guluronate content alginate carrier was effective for only 3 hours. This work demonstrates that injectable alginate solutions may be effective as local drug delivery vehicles. Lee et al. [29] recently developed an in situ gelling modified poly(guluronate) that cross-links by reaction between aldehyde and adipic acid dihydrazide groups. Primary rat

osteoblasts were subcutaneously injected into mice and formed mineralized tissue within 9 weeks of implementation of the gel matrix. Enzymatic cross-linking In situ formation catalyzed by natural enzymes has not been investigated widely but seems to have

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some advantages over chemical and photochemical approaches. For example, enzymatic processes operate efficiently under physiologic conditions without the need for potentially harmful chemicals such as monomers and initiators. Adjusting the amount of enzyme also provides a convenient mechanism for controlling the rate of gel formation, which allows the mixture to be injected before gel formation. Recombinant forms of human enzymes (e.g., tissue transglutaminase and factor XIII) that cross-link proteins in a highly specific manner have been engineered and can be used to induce in situ gelation of protein and peptide mimetic hydrogels. Although they perform with different substrate specificities, tissue transglutaminase and FXIII are calcium-dependent enzymes that catalyze covalent coupling between the side chain amine of lysine (amine donor) and the side chain amide of glutamine (amine acceptor) residues of proteins, which results in the formation of intermolecular cross-links. Sperinde and Griffith [45] used guinea pig liver transglutaminase to cross-link a glutaminamidemodified PEG and a poly(lysine-phenylalanine) copolymer into a network gel. Gel formation under the conditions of the study occurred over several hours, and the approach was found to be useful for entrapping cells within the hydrogel network. A follow-up study demonstrated that the architecture (extent of branching and number of reactive functional groups) of the cross-linkable macromer significantly affected the kinetics of the gelation reaction [44]. An increasing number of arms on the macromer led to a decrease in gelation time. Westhaus and Messersmith [55] used human recombinant FXIII and fibrinogen, a natural enzyme-substrate pair to form protein hydrogels rapidly (less than 10 minutes) at physiologic temperature, and introduced the concept of thermal control of enzyme-mediated gelation through the use of temperature-responsive liposomes that trigger conversion of the enzyme from inactive to active state by release of calcium. Chemical polymerization There have been numerous reports of chemical polymerization of monomers or macromers for in situ gel formation. Yaszemski et al. [56] have developed an injectable form of poly(propylene fumarate) that hardens in situ by a polymerization reaction between the polymer and N-vinyl pyrrolidone as cross-linker. This polymer is degradable and has been formulated with calcium phosphate particles to form a polymer composite that has potential for use in orthopedic applications [20,33 – 35]. A similar strategy has been

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reported by the same investigators to form hydrogels based on poly(propylene fumarate-co-ethylene glycol), whose polymerization was initiated by benzoyl peroxide [48 – 51]. Photopolymerization Photopolymerization is commonly used for in situ formation of biomaterials. A solution of monomer or reactive macromer and initiator can be injected into a tissue site and the application of electromagnetic radiation used to form the gel. Acrylate or similar polymerizable functional groups are typically used as the polymerizable groups on the individual monomer and macromers because they rapidly undergo photopolymerization in the presence of a suitable photoinitiator. Typically, long wavelength ultraviolet and visible wavelengths are used. Short wavelength ultraviolet is not used often because it has limited penetration of tissue and is biologically harmful. A ketone, such as 2,2-dimethoxy-2-phenyl acetophenone, is often used as the initiator for ultraviolet photopolymerization, whereas camphorquinone and ethyl eosin initiators are often used in visible light systems. These systems can be designed readily to be degraded by chemical or enzymatic processes or can be designed for longterm persistence in vivo. Sawhney et al. [41] synthesized various telechelic PEG-based macromers that consist of a linear central core of PEG flanked on either side by oligomers of degradable polyesters (PLA, PLGA, PGA) and terminated by a polymerizable acrylate group [41]. PEG was chosen for its biocompatibility and water solubility, and lactide and glycolide were chosen for their biodegradability. A high degree of variability of polymer properties, such as solubility and degradation rate, was achieved by tailoring the molecular weights and compositions of the degradable and nondegradable segments. Aqueous solutions of these macromers were rapidly photopolymerized to yield degradable hydrogels that have been investigated for use in prevention of postsurgical adhesions and for inhibition of thrombosis and intimal thickening at arterial surfaces damaged during angioplasty. Elisseeff et al. [10 – 12] have demonstrated the use of transdermal photopolymerization as a technique to induce hydrogel formation noninvasively. To demonstrate this technique, they used PEG-diacrylate along with a ketone as the initiator. In an in vivo experiment, a mixture of the polymer, initiator, and chondrocytes was injected subcutaneously into athymic rats and transdermally irradiated with visible

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light to induce gelation. Implants harvested after 2, 4, and 7 weeks were found to contain viable chondrocytes that produced a neocartilage-like matrix that contained collagen and proteoglycans. Glycosaminoglycan and collagen production increased with implantation time, whereas the chondrocyte morphology changed from rounded to elongated, more like fibroblasts. It was hypothesized that this morphologic change was caused by the ectopic implant placement and lack of mechanical stimulation. In a recent study, alginate and hyaluronan also have been modified with methacrylate compounds for use in visible light photopolymerization [43].

Self-assembling systems Self-assembly of proteins, peptides, lipids, and other biomolecules is emerging as an attractive approach to engineering stimulus-responsive materials with controlled nanostructures and microstructures for in situ forming biomaterials for medical and dental applications. Recent efforts to design such systems include proteins that self-assemble into hydrogels via leucine zipper motifs [36], cyclic peptides that self-assemble into nanotubes [6,16,18], and peptides that form fibrillar b-sheet networks [1,23,37,58]. Many of these systems self-assemble in response to an applied stimulus, such as a temperature or pH change, and may lead to useful new biomaterials for medical and dental applications. Among the self-assembling systems that can be exploited for in situ formation include proteins that contain blocks that aggregate under specific conditions, such as the leucine zipper motif [36]. Under appropriate conditions of pH and temperature, leucine zipper blocks assemble into dimeric and multimeric aggregates of helical segments; self-assembly is modulated by hydrophobic interhelical interactions and pH-dependent changes in charged residues along the leucine zipper block. Petka et al. [36] used recombinant DNA methods to engineer artificial proteins that consist of a central water soluble polyelectrolyte segment flanked by leucine zipper regions. The proteins undergo reversible gelation in response to changes in pH or temperature. For example, at low pH the helix-helix interactions that give rise to gelation are stabilized because the side chains on the glutamic acids are protonated, whereas at higher pH the deprotonation causes electrostatic repulsion of the helix-helix interactions, which results in disruption of the gel. Cappello et al. [5] have reported the use of proteins containing b-sheet forming silk- and elastin-like motifs for thermal gel formation. These

biosynthetic approaches are amenable to tailoring these systems for predetermined biological and physical properties of the gels. Some peptides with alternating hydrophobichydrophilic residues are known to self-assemble into b-sheet structures, often in an ionic-strength-dependent manner [3,57 – 59]. For example, the peptide H2N-(AEAEAKAK)2-COOH (EAK16-II) is known to produce b-sheet fibrillar gels in the presence of millimolar concentrations of monovalent salts [58,59]. b-sheet structure is favored by the strict alternating hydrophobic-hydrophilic primary structure of the peptide, which positions all hydrophobic side chains on one side of the b-sheet and all hydrophilic side chains on the other [53,54]. Saltinduced self-assembly of this peptide may be driven by the shielding of electrostatic repulsive forces with increasing ion concentrations, which allows attractive hydrophobic and van der Waals forces to dominate [4]. A derivative of this peptide has been used to study neurite growth in a rat model [21]. It may be possible in the future to exploit these peptides as injectable gel-forming biomaterials for drug delivery and tissue engineering.

Summary In this article the authors discussed various existing and emerging strategies for in situ formation of biomaterials. The benefits of in situ formation are that it is less invasive than conventional surgical administration, and the material takes the exact shape of the cavity or defect being filled. Most biomaterials used for in situ formation are hydrogels. The mechanisms by which in situ formation occur involve one or more of the following: a change in temperature or pH, physical changes in the biomaterial such as diffusion and swelling, direct chemical or photochemical reactions, and more recently, self-assembly of molecules. In situ forming biomaterials can be used for many types of medical and dental treatments and procedures, including drug delivery, wound healing, and tissue engineering. The use of in situ forming biomaterials for soft tissue regeneration, such as regeneration of the periodontium and cartilage tissue, are only beginning to be explored.

Acknowledgment The authors thank Dr. Franco Vittur for providing Fig. 3, and NIH grants DE13030, DE12599, and DE13076 for support.

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Surface modifications of implants Clark M. Stanford, DDS, PhD Dows Institute for Dental Research, N447 Dental Science Building, College of Dentistry, University of Iowa, Iowa City, IA 52242, USA

Since the early 1980s, the use of endosseous dental implants for the support of dental restorations has created a revolution in the routine approach to dental care. The high success rate for this elective procedure occurs through the initial stability that is provided by the amount, quality, and distribution of bone within the proposed implant site [111]. Typically, the integration of an implant is characterized by a series of clinical rules based on the concepts of engineering statics, implant surface technologies and conventional prosthodontic principles [123]. These relatively crude clinical parameters involve a lack of signs and symptoms of pathology, a lack of mobility, and a radiographic assessment of the interface [5,122]. The clinical success of implants relies on shortterm surgical issues (adequate bone volume, minimal surgical trauma) and long-term biologic issues of masticatory load-mediated bone adaptation [123]. This long-term response is only achieved through dynamic modeling (any net change in bone shape) and remodeling (the continuous turnover of bone without a net change in shape or size) processes of bone (which are initially woven in nature) [17,123]. This adaptive capacity allows bone to withstand the tolerances of clinical function (e.g., the accuracy of technical procedures, masticatory loading parameters) while creating a structural material capable of supporting clinical loads over long periods of time. Whereas high success rates hold for certain anatomic regions,

This work was supported in part by the Roy J. Carver Charitable Trust (Muscatine, Iowa) and the University of Iowa Centennial Fund. E-mail address: [email protected] (C.M. Stanford).

the bony response within the thin cortical plates and diminished cancellous bone that characterize type IV bone is considerably less successful with conventional machined surfaced implants (e.g., 65% – 85%) [13,55, 56,75,76,141]. In a meta-analysis, Lindh et al. [76] observed that implant success in the posterior maxillae is less than other regions of the mouth and depends highly on adequate bone volume in the area. For these reasons, the response of trabecular bone to the mechanical environment is a critical factor, especially in regions of the jaw, such as the edentulous posterior maxillae, where the cortical thickness and local material properties are insufficient to withstand occlusal forces. Strategies to increase the local quantity and quality of osseous tissue at the interface are an important means to provide predictable implant therapy for patients with poor bone quality, immediate loading protocols, and implant design strategies that potentially decrease the need for a large number of endosseous-style implants for the rehabilitation of a patient. To increase the predictability of implant therapy, significant efforts have gone into developing implant biomaterials that hold the promise of improving clinical success. Unfortunately, the dental implant market is currently being driven by the perceived need for biomaterial ‘‘innovations,’’ whereas sometimes the real need is a company’s patent and marketing position. It is important for the clinician to understand what the original objectives were in changing an implant surface that otherwise may have been functioning sufficiently for a long period of time. The purpose of this article is to provide a framework and background for the clinician to understand the relative merits of various ‘‘innovations’’ (in development or already marketed) to the

1042-3699/02/$ – see front matter D 2002, Elsevier Science (USA). All rights reserved. PII: S 1 0 4 2 - 3 6 9 9 ( 0 2 ) 0 0 0 1 6 - X

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surface topography of titanium endosseous-screw type implants.

Implant macroretentive features Implants used in the oral environment have one of three major types of macroretentive features: screw threads (tapped or self-tapping), solid body press-fit designs, and sintered bead technologies. Each of these approaches is designed to achieve initial implant stability and create large volumetric spaces for bone ingrowth. An important biologic principle of bone is that it responds favorably to compressive loading (without the presence of a periodontal ligament (PDL)) but not to shear forces [123]. Screw thread designs have been adapted to achieve a compressive loading of the surrounding cortical or cancellous bone. For instance, certain implant designs (ITI/Straumann, Institut Straumann AG, Waldenburg, Switzerland) use a 15
cutting thread profile, which creates primarily compressive versus shear interfacial stress. This thread profile has a rounded tip (reducing shear forces at the tip of the thread) that seems to maintain bone in the compressive zone beneath the thread profile [117 – 119]. With a desire to improve initial bone stability, various implant designs have incorporated dual (or more) cutting thread profiles (with or without a screw-based press-fit) in which two sets of threads cut at different relative locations in the osteotomy upon placement as a means to reduce initial stripping (e.g., upon overseating) and loss of primary stability (Mark III and Mark IV, Nobel Biocare, Go¨teborg, Sweden). Still other thread designs (Microthread Astra Tech AB, Mo¨lndal, Sweden) have focused on reducing the surrounding shear forces by reducing the height of the thread profile (reducing the contribution of any one thread) with an increase in the number of threads per unit area of the implant surface [45]. This approach has the added benefit of increasing the strength of the implant body by increasing the amount of remaining wall thickness of the implant body [12]. Finally, orthopedic prostheses (e.g., femoral stems, pelvic acetabular caps, knee prosthesis) have used various sintering technologies to create mesh or sintered beads as a surface for bone to grow into. The application of this technology to dental implants has involved attempts to improve the success rate of short implants (less than 10 mm in length), which are associated with the highest failure rates [76]. At the University of Toronto, sintered bead technology has led to the development of one commercially available

implant system (Endopore Innova Corp, Toronto, Canada) that shows favorable clinical results (93% – 100% relative success at 3 to 5 years) with short implants (7.7 mm) even in the posterior maxilla [37 – 40,79,98 – 101,129].

Implant microretentive features After placement of an implant into a surgical site, there is a cascade of molecular and cellular processes that provides for new bone growth and differentiation along the biomaterial surface. The goal of several current strategies is to provide an enhanced osseous stability through microsurface mediated events. These strategies can be divided into those that attempt to enhance the inmigration of new bone (e.g., osteoconduction) through changes in surface topography (e.g., surface ‘‘roughness’’), biologic means to manipulate the type of cells that grow onto the surface, and strategies to use the implant as a vehicle for local delivery of a bioactive coating (adhesion matrix or growth factor such as a bone morphogenetic protein (BMP)) that may achieve osteoinduction of new bone differentiation along the implant surface. What is surface roughness? One means to improve implant success is through methods to increase the amount of bone contact along the body of the implant. Although it may seem obvious that increased surface roughness of implants leads to greater success, it is not clear that just any ‘‘roughness’’ is advantageous. The term ‘‘roughness’’ itself can be ambiguous [137]. In implant design, it is usually assumed that a greater surface area (per unit of bulk metal surface) is an objective by various means to enhance the surface roughness of the implant surface. This enhanced surface area then allows a greater area for load transfer of bone against the implant surface [21,46,47,139]. Surface roughness is often a poorly described characteristic of implant surfaces that makes comparisons between implant systems difficult [137]. In evaluating an implant surface there are macroscopic and microscopic features that, when combined, are used to describe the surface ‘‘topography.’’ For instance, various screw thread profiles are used as macroscopic features (pitch, number, sharpness) combined with microscopic features (surface pits created by grit blasting) that may or may not be coupled with acid etching of the roughened surface. These micromechanical features influence the process of secondary integration (bone growth, turnover and remodeling)

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Fig. 1. Scanning electron microscopy (SEM) and atomic force microscopy (AFM) of a machined surface implant. Note the groove lines created during the milling process create local complex surface topography at different levels of the tip of the thread, along the side, and at the base of the individual groove (original magnification 1000; bar = 10 mm).

[123]. One advantage of acid etching is to increase the roughness of the already grit blasted surface and create a nanometer-scale topography that allows bone to grow into and maintain the implant surface under elevated shear forces (Figs. 1 – 4) [17,62]. Implant design features conventionally were believed to need surface pores or ‘‘pits’’ of 100 mm or more in diameter for ingrowth of bone, although clinically relevant surface roughness actually may be much finer (on the nanoscale level) [70]. Various titanium surfaces have used surface roughness created either through a grit blasting and etching procedure or blasting of the surface alone by using tightly controlled conditions to obtain a predefined optimal surface topography. One such optimization criterion has been proposed [47]. This criterion suggests that an implant surface has an optimal balance between pore size on the surface (pore sizes of 1 – 5 mm diameter and 1 – 5 mm in depth), which optimizes the shear strength of the

individual bone ingrowth into any one pit with the need to have as many ‘‘pits’’ on the surface as possible [46,47]. This theoretical model relates the relative shear strength of bone (ts) with a topography of interfacial load carrying capacity for a surface topographic feature (‘‘pit effectivity factor’’) combined with a value for the optimal number of pit per unit area of the surface (‘‘pit density factor’’) [47]. Using this optimization scheme (an unusual approach in the dental field but one that is highly desirable), a theoretically optimal surface has been used to help design a commercial implant system (Astra Tech AB TiO blast) [29]. These topographic features are then combined with macroscopic implant thread profiles (pitch, angle, and position) to provide a high compressive loading (low shear) implant interface [45]. This surface topography is currently available from one manufacturer (See Fig. 2) and shows promise by maintaining crestal bone at the level of the head of the implant [1,7,8,32,36,42,53,82,92].

Fig. 2. SEM and AFM of a TiO2 blasted implant surface (TiOblast, Astra Tech AB, Mo¨lndal, Sweden). Grit blasting with TiO2 creates multiple surface pits and roughened surface undulations that create an enhanced area for bone ingrowth. Note the lack of a repeated groove pattern (original magnification 1000; bar = 10 mm).

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Fig. 3. SEM and AFM of an acid etched implant surface (Osseotite, 3i/BioMet Corp, Palm Beach Gardens, FL). Acid etching with HCL/H2SO4 creates multiple surface tags on the implant surface, although these tags are not as deep as on the grit blasted surface (original magnification 1000; bar = 10 mm).

Recently, a refinement of Hansson’s grit-blasted topographic theory has been evaluated in a rat tibia model by a combined aluminum oxide grit blasting and hydrochloric acid (HCL) etching. This study demonstrated a 62% greater bone formation than a typical machined surface or a nonoptimal ‘‘rough’’ surface (blasting alone). Interestingly, the typical measurements used to characterize the surface roughness (Rp-v, Ra, and V) were not significantly different between the two blasted surfaces [2]. This discrepancy between descriptions of surface roughness and a biologic outcome (e.g., pullout strength) is not uncommon and illustrates the need to define carefully what is meant by surface roughness [137]. Studies of these TiO2 blasted surfaces from a histologic perspective are only recently available. In one such study, small TiO2 blasted (25 mm size) or machined surface microimplants were histomorphometrically compared in a human retrieval trial of 27 edentulous patients (two implants being placed in

each patient). The surfaces were evaluated for topographic features and bone contact and area filling of the threads evaluated from longitudinal sections of each microimplant. After 6 months of healing, significantly greater bone contact was observed with the blasted surface (40% versus 10% in either jaw) and greater bone fill within the threads was observed, especially in the mandible [54]. Surface roughness by blasting or etching Various studies also have addressed the issue of surface roughness through various means of grit blasting followed by a surface etching or coating procedure. This has included titanium plasma spray [21], abrasion (TiO2 blasting or use of soluble abrasives), combinations of blasting and etching (e.g., Al2O3 with H2SO4/HCL) [21], thin apatite coating [130], or sintered beads [40]. Commercially available roughened surfaces using the large grit blasted and

Fig. 4. SEM and AFM of a large grit blasted and acid etched implant surface (SLA, ITI/Straumann, Institut Straumann AG, Waldenburg, Switzerland). Grit blasting followed by etching of the surface creates a uniform repeating pattern of surface topography (original magnification 1000; bar = 10 mm).

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acid etched surface (ITI/Straumann’s SLA surface) have shown laboratory and clinical evidence of elevated success rates in areas of the posterior maxilla (Fig. 4) [18,19,21,22,28]. The SLA surfaces is one example of a commercially available surface technology that is derived from original animal studies (mini-pig) in which electropolished cpTi was compared to a range of roughened surfaces (some with etching) and a calcium phosphate coated surface hydroxyapatite (HA). This study documented that the SLA surface achieved equivalent bone contact (50% to 60%) as the HA surface but was able to achieve this without the need for a soluble coating [21]. Wong et al. [142] looked at press-fit implants in trabecular bone sites in the mini-pig model over 12 weeks and characterized a near linear increase in mechanical strength of the interface as function of increasing roughness. The role of the roughened surface is complex because the actual strength of bone contact against the titanium oxide surface is low (4MPa or less); weak enough that without the surface topography (e.g., electropolished surfaces) little bone contact occurs [17]. Clinically, the combination of large grit blasted and etched surfaces (ITI/Straumann SLA surface) in a one-stage surgical procedure, has documented more than 10-year cumulative survival rates of 96.2% (30.5% being in the maxilla) [20]. An alternative commercially surface etched implant design (see Fig. 3) uses a combination of HCL/ H2SO4 to create a surface topography (Osseotite, 3i Implant innovations/BioMet Company, Palm Beach Gardens, FL). This alternative roughened surface seems to have enhanced biomechanical strength in animal studies [68]. This surface technology currently is being evaluated in clinical trials (relative to machined surface implants). Data seem to support elevated success rates favorably in conventional two-stage protocols (delayed and early loading, for example, at 2 months after placement) [43,72,73]. The continued application of a roughened surface topography seems to have become the current state-of-theart with multiple systems coming out with new or modified products to reflect this increased interest. Biologic response Given the clinical interest in surface technology, it begs the question if biologic responses are seen with these roughened surfaces. In other words, are the favorable clinical responses simply a function of greater bone contact or can specific surface technologies influence the number of bone formative cells (osteoconduction), the type of cells present (osteoinduction), or the amount of bone cell expression and

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matrix expression per cell (osteopromotion)? Various studies have tried to address these issues using in vitro systems. For instance, the expression of matrixrelated proteins such as alkaline phosphatase (AP) and type I collagen were enhanced on SLA surfaces [35]. The mechanism by which topography influences osteoblast expression seems to be mediated by the protein kinase A and PL A2 pathway [14]. The topography also influences subsequent expression of osteoblast-mediated cytokines and growth factors. When osteosarcoma cells (MG-63) are grown on roughened surfaces, an increase in transforming growth factor-b and interleukin-1b expression is observed [10]. These responses are prostaglandin mediated and lead to a decreased proliferation on characterized rougher surfaces with an increase in cellular phenotypic markers of differentiation (AP activity, osteocalcin) [10]. Similar responses also have been described with primary cells derived from rat calvarial cells [77,78]. Many of these features have been described to act through a phospholipase A2 pathway that leads to a surface roughness-mediated increase in the expression of specific prostaglandins (e.g., PGE2) [121]. This suggests that osteoblasts have the capability to respond directly to cell shape changes induced by growth on a roughened surface or the matrix deposited and oriented on roughened surfaces [14]. As is typical in biomedical science, the story is more complex. Recent evidence indicates that the reduction in cell growth (and the corresponding increase in elaboration of an extracellular matrix) seems to occur through integrin-mediated signaling that is modulated by Vitamin D [14,78]. This hormone-based modulation of cell response to the surface topography seems to elicit cross-talk activation of two major signal transduction pathways (protein kinase A (PKA) and protein kinase C (PKC)) that converge on a central proliferation controlling pathway for cell growth, the mitogen-activated protein kinase pathway [121]. These in vitro studies emphasize the role of how cell shape influences osteoblast interactions with implant surfaces and the mineralizing extracellular matrix and only serves to highlight the complex ‘‘waltz’’ that goes on in vivo [30,31,83,116,126]. Surface chemistry and biologic laboratory studies with titanium alloy (Ti6AL4V) demonstrate altered oxide composition and altered biologic responses as a function of grit blasting and surface etching procedures [6,31]. Although these observations demonstrate cellular responses to how osteoblasts may interact with matrix absorbed onto the surface of an implant, they do not provide information on the role that the surface (and the resultant matrix topography)

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has on altering osteoblast cell shape and subsequent differentation (events critical to the ingrowth of new cells), the process of osteoconduction, and the possibility that the implant topography may lead to enhanced differentation of osteoblasts through alterations in cell shape and the consequent cell shape/ adhesion-mediated differentation on the implant surface. Whereas in vitro approaches are valuable, especially for understanding basis mechanisms underlying osseointegration (and serve as a valuable design criteria), clinical evaluation of surface modifications is important for evaluating the impact of these changes. Unfortunately, the rate of ‘‘innovations’’ in implant surface topography often outpaces any clinical evaluation before a modified design is introduced to the general marketplace. Alterations of the surface oxide Various attempts have been made to alter systematically the surface topography and composition of the oxide surface on commercially pure and Ti-6Al4V titanium alloy. These approaches have involved various techniques to increase the surface oxide thickness, create repeated patterns (e.g., through additive or subtractive lithography), or control oxide composition through the sterilization techniques used during or after manufacturing. Titanium, being a nonnoble metal, spontaneously forms a 3- to 5-nm thick oxide surface (primarily dioxide and trioxides) when exposed to air [124,136]. In vitro studies have been performed to characterize changes in the cell and molecular regulation of bone-related matrix proteins as a function of the oxide composition that results from sterilization procedures [61,64,65,71, 88,125]. These studies emphasize the need for implants to be manufactured, packaged, and clinically handled in a careful manner [66,67,95]. Can the results of a simple cleaning and sterilization process impact subsequent biologic healing? This has been a significant debate for several years and has led to significant regulatory discussions in the European and North American markets. Most manufactures provide their implant systems in a preencapsulated form to minimize contamination. Recent studies on surface sterilization have led to implant surfaces having a relatively enriched oxide content of fluoride. Fluoride seems to play a significant role not only in modulating the physical chemistry of hydroxyapatite formation but also in stimulating new bone growth, matrix expression, and subsequent mineralization [11,41,57,74,86,87,89 – 91]. In vitro and animal studies have demonstrated enhanced trabecular bone growth along the sides of the fluoride-

containing roughed implant surface that passes through the medullary cavity (relative to equivalently rough but nontreated implants). In a study using Sprague-Dawley rat tibiae, two different threaded implants (1.5  2mm) were compared that were either TiO2 grit blasted or blasted and treated with a fluoride containing solution. After 21 days, a significantly greater amount of bone contact had occurred on the fluoride coated surfaces (55.5% versus 34.2%; P < 0.027) [3]. Histomorphometrically, the flouride-exposed surface demonstrated greater osteoconductive growth of trabecular bone in the marrow cavity. In a rabbit tibial model, Wennerberg and Albrektsson [137,140] demonstrated enhanced pullout strengths (57% increase from 54 to 85 ncm at 3 months) and a more rapid formation of bone contact on implant surfaces that were exposed to this fluoride surface treatment. This surface conditioning is currently undergoing human clinical trials. These studies serve to emphasize how relatively minor changes to surface technology can create the potential for significant changes in bone contact and biomechanical strength that have the potential to allow more rapid clinical function of implants. One method to increase overall surface roughness has been to alter the surface oxide that forms on the implant surface. During the preparation of a conventional machined surface implant, a titanium oxide surface forms when exposed to air, which results in a 3- to 5-nm thick oxide composed of TiO2 and variant forms of TiO3. Various case reports of retrieval material have suggested that the oxide surface enlarges as a function of time once an implant is in contact with bone and results in a growth from 5 nm at placement to approximately 200 nm after 5 years [85]. Sundgen et al. [127] evaluated the composition of this enlarged oxide surface and determined that it contained calcium and phosphate within the oxide surface which suggests that part of the integration process was a mutual growth of the oxide surface and bone growth to the surface of the implant. These early observations led to a series of developmental projects to create an enlarged oxide surface through various thermal and electrochemical means [33,67,69,70,80,138]. Larsson et al. [69] evaluated machined and electropolished titanium samples with and without an enlarged anodically formed surface oxide layer. In this study, implants were evaluated over 6 weeks in a rabbit cortical bone model. Surfaces of similar roughness values (as measured by conventional surface analytical techniques) demonstrated significantly greater bone contact on the enlarged oxide layers [69].

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Recently, one manufacturer introduced an implant design with a microporous thickened oxide layer (1 – 10 mm thick) created through an electrochemical process (TiUnite, Nobel Biocare, Go¨teborg, Sweden). This oxide surface has an open porous structure with various random ‘‘pits’’ or holes of variable dimension (1 – 5 mm in diameter) that create a roughened surface that seems to allow greater bone contact [44]. Early animal studies using greyhound dog mandibles have indicated that this polycrystalline oxide surface has greater (twofold) removal torque than an equivalent machined surface after 10 weeks of healing [50]. Studies in New Zealand white rabbits in the femur/ tibial region suggest that the enhanced oxide surface elicits greater bone-to-implant contact and higher removal torques [4]. It remains to be seen how this implant design will fare in the human clinical trials currently underway.

Biologic modifications of implant surfaces The biologic modifications of implant surfaces through addition of biologically active macromolecules are a potentially exciting new area of implant tissue engineering. These emerging technologies use the understanding of how the initial coagulum (especially platelets and macrophages) and subsequent osteoblasts interact with the extracellular matrix (e.g., through integrin receptors). The initial matrix that is deposited on the implant surface then modulates subsequent cellular differentiation and may even influence the role of specific regulatory growth factors (e.g., the BMP family) [17,123, 124,126]. The integrins are a heterogeneous family of transmembrane receptors that link the extracellular matrix to the internal cytoskeleton. Through binding to the extracellular matrix (ECM) proteins they mediate intracellular adhesion-mediated signal transduction pathways that contribute significantly to the development of tissues and organs, wound healing, cell motility, cell growth and differentiation, and programmed cell death (also known as apoptosis) [34, 58,63,113 – 115, 120,143]. It is important to emphasize that an implant surface (no matter what its surface chemistry) must interact with the initial serum-based coagulum at the time of placement. Unfortunately, although the initial interaction of blood with an implant is logically important, little is actually known about this. In vitro studies have suggested that an etched roughened surface leads to greater accumulation of platelets and acts as a potential source of platelet derived growth factor (PDGF) and other mitogenic factors than a machined surface [93,97].

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Kanagaraja et al. [59] and Nygren et al. [93] used a series of different immunochemistry approaches to evaluate fibrin and fibrin/platelet aggregation with an implant surface [52,135]. These studies indicate that an elevated roughness leads to enhanced platelet activation and alterations of the fibrin scaffold. Subsequent to the fibrin clot formation, a complex interaction occurs among macrophages, the scaffold, and undifferentiated mesenchymal stem cells. These stem cells are a viable source of new bone forming cells that proliferate in a marrow chamber under the complex regulatory control of the macrophage population of cells in the healing osteotomy [9,51,81, 96,112]. One future tissue engineering approach of potential value is to modulate this regulatory control to amplify the differentiation of marrow-derived stem cells though biomanipulation of the implant surface. The manner in which the fibrin scaffold is manipulated is one key to the future of implant surface technology. A recent approach to modifying the titanium oxide surface uses placement of various configurations of the commonly known recognition sequence for integrins (a tripeptide sequence of ArgGly-Asp or RGD) on the surface of an implant surface. This RGD sequence mediates cell adhesion and is present in several extracellular matrix proteins (e.g., fibrin, collagen, fibronectin, vitronectin, osteopontin, and, importantly, bone sialoprotein). Many, if not most, mesenchymal cells possess integrin receptors, and the adhesion to an RGD-coated surface may be nonspecific, but several groups are attempting to regulate the type of cell adhesion that occurs by modulating the sequence of proteins in the linker region (the region of protein that attaches the RGD sequence to the metal substrate) and exploring various means to attach adhesion sequences (e.g., repeated regions of RGD sequence) through covalent attachment to the surface of the implant [17,48,49,84,102, 103,105 – 110,128]. Recent work has focused on determining the optimal density of RGD adhesion peptides on an experimental implant surface to elicit osteoblast growth and differentiation [106]. Osteoblastic cells interact with the matrix through RGD-dependent pathways and with the surface contours of the surface. This property of ‘‘contact guidance’’ is an important part of understanding the behavior of how preosteoblasts interact with the complex fibrin/platelet scaffold. Several elegant in vitro and in vivo studies document the importance of not only the roughness of an implant surface but also the effect of the direction of the pattern (or epitaxis) [15,16,23 – 27,94,104,131 – 134]. These studies demonstrate that fibroblasts and osteoblasts are capable of recognizing repeated surface features (e.g., lines,

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grooves, and other repeated defects created in milling). Cells seem to align and grow in a directional pattern directed by the surface features of the substratum [15]. Highly repeated, nanoscale surface features are capable of being formed with inexpensive photolithographic approaches. These repeating surface features are capable of combining properties of surface chemistry (e.g., surface energy) with biologic attachment of adhesion/matrix peptides [15,60]. Similar photolithographic approaches have been used to create repeated patterns to generate controlled alternating domains of N-(2-aminoethyl)-3-aminopropyl-trimethoxysilane and dimethyldichlorosilane as a means to control the adsorption of naturally occuring RGD adhesive proteins in serum (especially vitronectin) [84,108]. In this way, a bioactive surface may be generated that uses the natural adhesive proteins in the blood plasma at the time of implant placement. Multiple issues of bioavailability and biologic stability of these peptides still are being worked out, but promise holds with such techniques. The covalent coupling of bioactive peptides also includes the possibility of coupling growth factors to the implant surface. Over the past 30 years various growth factors have been described with various growth promoting (mitogenic) and inductive (differentiation) activities. One large class of such regulatory macromolecules, the transforming growth factor-b – related bone morphogenetic factors (reviewed in another article in this issue), has the potential to play key regulatory roles as activators of other downstream growth factors, leading to an amplification of their apparent activity. This is an advantage of these factors and a potential worry because the cascade of events that are elicited by these factors must be controlled tightly. Uncontrolled and excessively high levels of bioactive peptides (as are commonly needed to elicit the putative bone inductive effects) lead to the possibility that implant surfaces that act as a local delivery system may not have an entirely beneficial effect.

Summary Implant dentistry has enjoyed an exploding number of new applications, techniques, materials, and devices since the early 1980s. In part, the acceptance of implant therapy has come about through the high success rates of the conventional surgical and restorative protocols. In an effort to increase the predictability of this therapy, especially with the emphasis on early and immediate loading, there has been an ongoing evolution in implant surface technologies and biomechanics. The market has seen the accept-

ance of some form of a roughened surface implant being sold by every major manufacturer. The future applications of surface technologies will see the continuing evolution to a biologic manipulation of bone and soft tissue chemistry for the long-term dental health of the patient.

Acknowledgment The author would like to thank Mr. Brad Kellogg for preparation of the SEM and AFM images.

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Recent advances in cosmetic materials Noah A. Sandler, DMD, MD Department of Oral and Maxillofacial Surgery, University of Minnesota, 7-174 Moos Tower, 515 Delaware Street SE, Minneapolis, MN 55455, USA

Emerging technology is allowing the permanent replacement of tissues for reconstructive and cosmetic purposes with minimal additional morbidity to the patient. By using genetic and biomaterial advances, many of these new techniques are minimizing the immunologic and foreign body responses seen in previous alloplast and autologous techniques.

Soft tissue Soft tissue replacement traditionally has been performed through the rotation of local flaps or the grafting of autologous tissue such as fat or dermis. Drawbacks of these procedures include the loss of tissue volume (especially with fat injections) [34] and donor site morbidity. In an attempt to reduce morbidity, several exogenous materials have been used. An ideal material for soft tissue augmentation has the properties of being safe and effective, is easy to administer, carries a minimal risk for infection, extrusion, or migration, produces a minimal inflammatory reaction, lasts for an acceptable length of time, and is cost effective [14]. An early material used in an attempt to reduce soft tissue defects, type II bovine collagen, is rendered ‘‘nonimmunogenic’’ by cleaving the end-protein fragments from the central helical portions of the molecule. This process, however, likely destabilizes the molecule, which causes unraveling of the helix and exposes other previously concealed aspects of the molecule. This process likely results in the 3% to 10% allergy demonstrated with the use of this material [2,9,13]. Resorption also is often seen with the need for retreatment every 2 to 6 months [7].

E-mail address: [email protected] (N.A. Sandler).

Alloplast materials have been used in an attempt to maintain a more permanent result. Bioplastique (Uroplasty BV, The Netherlands), Artecoll (Rofil Medical International, Breda, The Netherlands), and Gore-Tex (WL Gore, Phoenix, AZ) threads are biologically inert, nonallergenic materials. Artecoll and Bioplastique are mainly used in Europe, whereas Gore-Tex threads are Food and Drug Administration-approved. Bioplastique consists of polymerized silicone particles 100 to 600 mm in diameter in a gel carrier. The gel component is phagocytized and replaced with host collagen, which stabilizes the material. A drawback of the material is that palpable nodules can develop despite correct surgical placement [13,17,28]. Artecoll consists of 30 to 40 mm micron polymethylmethacrylate microspheres suspended in bovine collagen. Even when injected subdermally, as recommended, blanching and palpable thickening can occur [17]. Gore-Tex was developed in the 1960s and consists of polytetrafluoroethylene in nodules and fibrils with a microporous structure. Among its other uses, it has been used for facial soft tissue augmentation with good results [23,29]. This material is not incorporated into the tissue bed and can result in palpability, capsule formation, and distortion [11]. Infection and foreign body granulomatous reactions also have been reported [19]. In a review of 200 clinical studies reporting results of a series of implanted biomaterials of the face, polymer and ceramic materials had an overall infection rate of 3% and an exposure/extrusion rate of 1.2%; 4.6% of implants were removed because of implantrelated complications [16]. Bacteria, when they come in contact with the surface of an implant, secrete a layered polysaccharide matrix. The combination of bacteria and matrix is referred to as biofilm. Biofilm helps protect the bacteria from eradication by antibiotics. Slow growth of

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the biofilm can cause the development of a thickened capsule or gradual degradation of the implant material. The bacteria over time may become more dormant but may be reactivated by trauma or inflammation. This characteristic may account for the delayed development of infection around some implants [10,16]. Primarily because of large pore size, the newer implants allow a greater degree of tissue ingrowth with a greater potential to eradicate bacteria by normal surveillance and inflammatory mechanisms. The greater the ingrowth of biologic tissue, the greater the potential stability of the implant [16]. In an attempt to reduce the immunogenic problems encountered with using a xenogenic material, a process to extract intact human collagen from skin was developed. Excised skin secondary to esthetic surgical procedures (i.e., rhytidectomy, abdominoplasty) is sent on ice to the processing unit (Collagenesis, Beverly, MA). A 5% collagen solution is then sent back to the physician to use on the same patient. On average, 2 in2 of skin is required to produce 1 mL of 5% Autologen. This material can be banked by the processing plant and stored for up to 6 months [14]. Dermalogen (Collagenesis, Beverly, MA) was developed in an attempt to eliminate the requirement for patients to donate their own skin and still provide human collagen tissue for injection. The material is derived from skin donors from approved tissue banks. The material is decellularized and aseptically processed into a collagen fiber suspension, which is suitable for immediate injection. All donors are screened through a detailed medical and social his-

tory for hepatitis B, C, HIV, HTLV-1, and syphilis. Skin samples also are tested for bacterial, fungal, or viral pathogens. Skin testing before placement is recommended, and 0.1 mL injections are supplied. These are injected in the volar aspect of the forearm, and signs of induration are assessed at 72 hours. Intradermal injection and overcorrection of 20% to 30% is recommended when injecting either Dermalogen or Autologen (Fig. 1) [14]. Posttraumatic wounds (e.g., wounds secondary to burns) frequently involve the loss of epidermis and dermis. Skin grafts traditionally have been used in these cases as an occlusive dressing, as a skin replacement, and as a stimulus for healing [12,21]. Recently, a tissue-engineered human skin product, Apligraft (Organogenesis Inc, Canton, MA), a bilayared skin equivalent, has been developed. This product consists of a bovine collagen fibroblast-containing matrix integrated on a sheet of stratified human epithelium. The fibroblasts and keratinocytes contained in this tissue are taken from screened neonatal foreskin serially grown in tissue culture, and they are viable, reproducing cells. By excluding Langerhans’ cells, this materials seems to be immunologically inert. With time, cultured allograft keratinocytes are replaced by host cells [21]. Even when the graft fails, significant healing takes place because growth factors and cytokines stimulate rapid healing from the wound margins or appendage structures within the wound and stimulate production of granulation tissue within the wound bed. These factors include interleukin 1,3,6,8 transforming growth factor-a, and plateletderived growth factor [18,21].

Fig. 1. (A) Preoperative view of ‘‘crow’s feet’’ region of a patient to be treated with Dermalogen injection. (B) Four month follow-up injection. Note decrease depth of wrinkles in the lateral canthal region. (Courtesy of Collagenesis, Inc, Beverly, MA.)

N.A. Sandler / Oral Maxillofacial Surg Clin N Am 14 (2002) 53–59

Tissue engineering also has led to the development of a living, metabolically active, inert dermal tissue, Dermagraft (Advanced Tissue Sciences, LaJolla, CA) [12]. Screened fibroblasts are similarly obtained form neonatal foreskin. Cells are cultured for 2 to 3 weeks. As the cells proliferate, they secrete dermal collagen, glycosaminoglycans, growth factors, and support proteins to produce a dermal matrix. Because the cells themselves secrete proteins, collagen fibers are arranged in parallel bundles with periodicity similar to normal tissue [31]. Clinical trials have demonstrated complete and rapid healing when this material was used [21]. AlloDerm (Lite Cell Corp., Brandsburg, NJ) is a an allogeneic, acellular dermal graft processed from tissue-banked skin. Similar to Dermagraft, this material is rendered nonimmunogenic secondary to the absence of cells [20,24]. The material is treated with an antiviral agent that has been demonstrated to inactivate HIV. There have been no reported cases of viral disease transmission in any patient treated with the material since 1992 [38]. This porous, dermal matrix allows ingrowth and colonization by host fibroblasts and endothelial cells, which allows for integration [7,23]. In addition to the advantages of lack of donor site morbidity and reduced procedure time, this material has been noted to maintain tissue volume better than autologous materials such as fat [7,8,26]. It has been used for cosmetic augmentation of the nasal dorsum, scar revision, glabellar contouring, and most recently, lip augmentation (Figs. 2 and 3) [20,38]. Tissue integration occurs as a result of neovascularization, which seems to be facilitated by vascular channels in the graft remaining intact [38]. The progressive growth

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of new fibroblasts has been shown to occur within 2 weeks of implantation, with the replacement by host collagen demonstrated at 5 to 8 weeks [38 – 40].

Bone Specifications for the ideal implant material differ from soft tissue replacement. The material should not be modified physically by surrounding soft tissue, should not induce inflammation or a foreign body reaction, does not produce allergy or hypersensitivity, is chemically inert, is noncarcinogenic, can resist mechanical strain, can be fabricated to various forms to fit the cosmetic defect, and can be sterilized [25,33]. Polymers consist of long chains of repeating subunits. The properties of the material are a result of the subunit and the type and number of cross-linkages. Silastic, a silicone polymer, has been used extensively in cosmetic augmentation as an onlay material. Because of its smooth surface, it can be inserted easily through small incisions. Typically, a fibrous capsule forms around the silastic implant, which helps to fixate the material. Recent studies have implicated the chemical degradation products of silastic breakdown in the destruction of the underlying osseous bed [15]. In an attempt to improve stability and reduce underlying destruction, porous implants were developed. A minimum pore size of 100 mm is necessary to allow bony ingrowth. Medpore (Porex Surgical, Fairburn, GA), a polymer that consists of high-density polyethylene, contains pores that vary in size from 125 to 250 mm capable of bony ingrowth. This material has been used for genial and malar augmentations

Fig. 2. Preoperative view of a patient with deep scarring with soft tissue defect lateral to the left commisure.

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N.A. Sandler / Oral Maxillofacial Surg Clin N Am 14 (2002) 53–59

Fig. 3. (A) Intraoperative view of sutures after Alloderm (LiteCell Corp., Branchburg, NJ) graft placement through an intradermal tunneling technique. (B) Postoperative view of commisure region with softening of the defect.

and in reconstruction of the external ear with good cosmetic results and minimal reported complications [25]. Because of the porous nature of these materials, there is an increased risk of infection; the manufacturer recommends antibiotic impregnation before implantation (Fig. 4). Hard tissue replacement (HTR) is a composite that consists of polymethylmethacrylate combined with polyhydroxyl methacrylate and calcium hydroxide coating. Polymethylmethacrylate has been used extensively in craniofacial reconstruction. Once polymerized, it is a strong and biocompatible material with minimal tissue response. The polymerized material is noncarcinogenic and radiolucent and possesses low

thermal conductivity. Polyhydroxyl methacrylate and calcium hydroxide allow the HTR implant to be markedly hydrophilic and porous and yet have substantial compressive strength. The calcium hydroxide also imparts a negative surface charge to the material, which may promote bone ingrowth and bonding [22]. HTR recently has been used in many of the same applications as Medpore, including malar and genial implants, with good results. Success of bone graft materials depends on the early establishment of osteoprogenitor cells, vascularity, and attachment of the newly formed osteoblasts. Whereas pore size of present materials has been noted to facilitate vascularization best at 150 to 500 mm,

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Fig. 4. Intraoral placement of Medpore (Porex Surgical, Fairburn, GA) malar augmentation.

limitations to bone development result in regions more distal to the blood supply forming cartilage rather than bone. The establishment of new vessels also depends on stability of the graft. In an attempt to provide stability during the healing phase, much research is being conducted in the study of biologic carriers [4]. Polymer scaffolds typically use polylactide, polyglycolide, or the copolymer, poly(lactic-gycolic) acid (PLGA). A small acute immune response has been noted by PLGA scaffolds. This immune response resolves when the material resorbs by hydrolysis [4,30]. Even with small pore size PLGA scaffolds (less than 50 mm pores), bone formation occurs throughout the scaffold rather than healing by creeping substitution. It is hypothesized that this process occurs secondary to a sponge effect, wicking tissue fluid and blood through the pores and stabilizing the hematoma [4,42]. Bioactive molecules and cells are added to these polymers in an attempt to enhance the osteoinductive capacity of cell recruitment, attachment, proliferation, and differentiation. Recently, bone morphogenetic proteins have been identified and their genes cloned. These molecules are normally synthesized and released locally. Scaffolds that incorporate bone morphogenetic proteins have demonstrated rapid bolus release, followed by continuous release as the polymer degrades [1]. Bone morphogenetic proteins added directly to demineralized freeze-dried bone have demonstrated an added beneficial effect in bone graft incorporation [36]. Other growth factors from the hematoma have been isolated and are being studied for their potential

to enhance bone healing. Fibroblast growth factor-2 enhances vascularization; when used with mineralized collagen, healing of segmental defects is improved [32]. Platelet-derived growth factor stimulates mesenchymal cell proliferation, which increases the number of potential osteoprogenitor cells [6]. It has been used effectively in bone healing when used with insulin-like growth factor-1 in periodontal defects [18]. Transforming growth factor-b1 stimulates cell proliferation and early differentiation in osteoblasts and chondrocytes [3,35]. The expression of this factor has been demonstrated to be upregulated at the time of active distraction during distraction osteogenesis. Decreased transforming growth factor-b levels have been noted during the period of bone consolidation and remodeling [16,27]. In contrast, the extracellular matrix proteins, fibronectin, osteocalcin, and collagen I fell below baseline during active distraction and were upregulated during consolidation. These studies demonstrate that cytokines that regulate osteogenesis may be upregulated before the upregulation of extracellular matrix proteins associated with bone formation [16]. Osteoblasts also respond to cytokines, including interleukin-1, prostaglandin E2, and tumor necrosis factor-a [16]. Gene transfer to the periosteum is a process in which DNA or RNA molecules are added into cells, which leads to the enhancement or inhibition of a preexisting function. Viral vectors, chemically mediated transfer, injection of DNA, and particle-mediated transfer techniques have been used. Currently, continued research is being conducted with a goal of

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enhancing bone formation within osseous defects using these methods to enhance bone production [16].

Cartilage Often, difficulty is encountered in reconstructing cartilaginous structures in establishing correct anatomic form. Rib grafts can be difficult to carve and contour in the intricate shape of an anatomically correct ear or nasal dorsum. Tissue engineering involves the morphogenesis of new tissue from dissociated cells and biodegradable polymers. In this manner, precise shapes can be used to reflect more accurately the correct anatomic form. Recently, a template in the shape of a human ear using a polyglycolide-polylactide polymer was implanted into a subcutaneous pocket in the dorsum of athymic mice. This structure was well maintained for 12 weeks after implantation, with new cartilage formation evident [5]. Fibrin monomers have been used since for the encapsulation of chondrocytes. This process creates a moldable gel with actively proliferating cells, which results in the production of a well-formed cartilaginous matrix [37]. More recently, a tissue-engineered adult human mandibular condyle was created using a similar polyglycolide-polylactide scaffold and copolymer hydrogel to suspend the chondrocytes [41].

Summary Recent advances in biomaterial engineering have created potential applications in cosmetic and reconstructive surgery. Many of these technologies may replace the current practice of autologous tissue transfer and eliminate the potential for donor site morbidity. Additional studies to evaluate the comparative esthetics and long-term stability of these materials are needed.

Acknowledgment The author wishes to thank Collagenesis, Inc for the images of patients treated with Dermalogen.

References [1] Agarwal CM, Best J, Heckman JD. Protein release kinetics of a biodegradable implant for fracture nonunions. Biomaterials 1995;16:1255.

[2] Barr RJ, Stegman SJ. Delayed skin test reaction to injectable collagen implant (Zyderm). J Am Acad Dermatol 1984;10:652 – 8. [3] Bonewald LF, Kester MB, Schwartz Z, Swain LD, Khare A, Johnson TL, et al. Effects of combining transforming growth factor beta and 1,25-dihydroxyvitamin D3 on differentiation of a human osteosarcoma (MG-63). J Biol Chem 1992;267:8943 – 9. [4] Boyan BD, Lohmann CH, Romero J, Schwartz Z. Bone and cartilage tissue engineering. Clin Plast Surg 1999; 26:629 – 45. [5] Cao Y, Vacanti JP, Paige KT, Upton J, Vacanti CA. Transplantation of chondrocytes utilizing a polymercell construct to produce tissue-engineered cartilage in the shape of a human ear. Plast Reconstr Surg 1997;100:297 – 302. [6] Cassiede P, Dennis JE, Ma F, Caplain A. Osteochondral potential of marrow mesenchymal progenitor cells exposed to TGF-B1 or PDGF-BB as assayed in vivo and in vitro. J Bone Miner Res 1996;11:1264 – 73. [7] Castor S, To W, Papay F. Lip augmentation with Alloderm acellular allogenic dermal graft and fat autograft: a comparison with autologous fat injection alone. Aesthetic Plast Surg 1999;23:218 – 23. [8] Chajchir A. Fat injection: long term follow up. Aesthetic Plast Surg 1996;20:291 – 6. [9] Charriere G, Bejot M, Schnitzler L, Ville G, Hartmann DJ. Reactions to bovine collagen implant. J Am Acad Dermatol 1989;21:1203 – 8. [10] Costerton JW, Lewandowski Z, Caldwell DE, Korber DR, Lappin-Scott HM. Microbial biofilms. Annu Rev Microbiol 1995;49:711 – 45. [11] Courtiss EH, Glicksman CA. Permanent lip augmentation employing polytetrafluoroethylene grafts. [discussion] Plast Reconstr Surg 1992;90:1091. [12] Eaglestein W, Falanga V. Tissue engineering and the development of Apligraf, a human skin equivalent. Advances in Wound Care 1998;11(14 suppl):1 – 8. [13] Eppley BL, Sidner RA, Sadove AM. Adequate preclinical testing of Bioplastique injectable material? Plast Reconstr Surg 1992;89:157 – 8. [14] Fagien S. Facial soft-tissue augmentation with injectable autologous and allogeneic human tissue collagen matrix (Autologen and Dermalogen). Plast Reconstr Surg 2000;105:362 – 73. [15] Freed JB. The increasing recognition of medullary lysis, cortical osteophytic proliferation, and fragmentation of implanted silicone polymer implants. J Foot Ankle Surg 1993;32:171 – 9. [16] Gosain AK, Persing JA. Biomaterials in the face: benefits and risks. J Craniomaxillofac Surg 1999;10: 404 – 14. [17] Hoffman C, Schuller-Petrovic S, Soyer HP, Kerl H. Adverse reactions after cosmetic lip augmentation with permanent biologically inert implant materials. J Am Acad Dermatol 1999;40:100 – 2. [18] Howell TH, Fiorellini JP, Paquette DW, Offenbacher S, Grannoble WV, Lynch SE. A phase I/II clinical trial to evaluate a combination of recombinant human platelet-

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[32] Radomsky ML, Thompson AY, Spiro RC, Poser JW. Potential role of fibroblast growth factor in enhancement of fracture healing. Clin Orthop Rel Res 1998; 355(suppl.):283 – 93. [33] Scales JT. Discussion on metals and synthetic materials in relation to soft tissue: tissue reaction to synthetic materials. Proc R Soc Med 1953;45:647. [34] Schuller-Petrovic S. Improving the aesthetic aspect of soft tissue defects on the face using autologous fat transplantation. Fac Plast Surg 1997;13:119 – 24. [35] Schwartz Z, Bonewald LF, Caufield K, Brooks B, Boyan BD. Direct effects of transforming growth factor-B on chondrocytes are modulated by vitamin D metabolites in a cell maturation-specific manner. Endocrinology 1993;132:1544 – 52. [36] Schwartz Z, Somers A, Mellonig JT, Carnes DL Jr, Wozney JM, Dean DD, et al. Addition of human recombinant BMP-2 to inactive commercial demineralized freeze dried bone allograft makes it an effective composite bone inductive implant material. J Periodontol 1998;69:1337 – 45. [37] Sims CD, Butler PE, Cao YL, Cassanova R, Randolph MA, Black A. Tissue engineered neocartilage using plasma derived polymer substrates and chondrocytes. Plast Reconstr Surg 1998;101:1580 – 5. [38] Tobin HA, Karas ND. Lip augmentation using an alloderm graft. J Oral Maxillofac Surg 1998;56:722 – 7. [39] Wainwright DJ. Use of an acellular dermal matrix (Alloderm) in the augmentation of full-thickness burns. Burns 1995;21:243 – 8. [40] Wainwright DJ, Madden M, Luterman A, Hunt J, Monafo W, Heimbach D. Clinical evaluation of an acellular allograft dermal matrix in full-thickness burns. J Burn Care Rehabil 1997;17:124 – 36. [41] Weng Y, Cao Y, Silva CA, Vacanti MP, Vacanti CA. Tissue-engineered composites of bone and cartilage for mandible condylar reconstruction. J Oral Maxillofac Surg 2001;59:185 – 90. [42] Whang K, Healy KE, Elenz DR. Emgineering bone regeneration with bioabsorbable scaffolds with novel architecture. Tissue Eng 1999;5:35.

Oral Maxillofacial Surg Clin N Am 14 (2002) 61 – 71

Skin and oral mucosal substitutes Kenji Izumi, DDS, PhDa, Stephen E. Feinberg, DDS, MS, PhDb,* a

Department of Tissue Regeneration and Reconstruction, Division of Reconstructive Surgery for Oral and Maxillofacial Region, Niigata University Graduate School for Medical and Dental Sciences, Niigata City, Niigata, Japan b Department of Oral and Maxillofacial Surgery, University of Michigan Medical Center, B1 – A 235 UH, Box 0018, 1500 East Medical Center Drive, Ann Arbor, MI 48109, USA

Biologic substitutes of human skin and mucosa have several prospective applications, including, but not limited to, (1) models of skin and mucosal biology and pathology, (2) treatment and closure of skin and mucosal wounds, (3) alternatives to animals for safety testing of consumer products, and (4) delivery and expression of transfected genes [3]. This article introduces the reader to the types of skin and mucosa substitutes that have been and are being developed in the area of tissue engineering for use in procedures for trauma, ablative oncologic resections, and reconstructive surgery. Skin and mucosal substitutes have a common set of requirements for the duplication of anatomic structures and physiologic functions that they are to emulate. For use in wound closure, the first definitive requirement is re-establishment of the epidermal barrier to fluid loss and microorganisms and alleviation of pain and enhancement of wound healing. In full-thickness skin loss, replacement of the epidermis and dermis is the preferred approach. Replacement of these tissue components also must minimize scar formation and restore acceptable function and cosmesis. A major advantage of the use of substitutes in wound closure is reduction or elimination of the donor site for skin and mucosal grafts. Success in the

This work was supported by a Grant-in-Aid for Scientific Research (No. 12771216) from the Ministry of Education, Science and Culture, Japan (KI) and from Grant No. DE13417 from the National Institute of Health, USA (SEF). * Corresponding author. E-mail address: [email protected] (S.E. Feinberg).

elimination or minimization of donor site morbidity could shorten recovery time and reduce the length of operative procedures. Approaches that have been used in the fabrication, manufacturing, and ‘‘tissue engineering’’ of skin and mucosa substitutes can be classified as (1) in vitro culturing of autologous and allogeneic keratinocytes, (2) in vitro tissue engineering of dermis composed of either artificial (collagen, glycosaminoglycans, polymers of polyglycolic, and polylactic acid) or allogeneic and acellular dermis, and (3) a bilayer of skin mucosa from a combination of (1) and (2). Three essential components are known to be necessary to engineer human skin and mucosa: cells, an extracellular matrix, and cytokines [30,42].

Skin substitutes A major milestone in the development of skin substitutes was the introduction of the in vitro technique of Rheinwald and Green [47] that involved the culturing of human keratinocytes into epithelial sheets suitable for autografts. These investigators used a combination of hydrocortisone, epidermal growth factor, and irradiated murine 3T3 fibroblasts to support the proliferation of keratinocytes on plastic substrates. The following ingredients were added to improve the culture media and to facilitate keratinoctye sheet formation: (1) insulin, to promote the uptake of glucose and amino acids; (2) transferrin, to detoxify iron; (3) hydrocortisone, to promote the attachment of cells and cell proliferation; (4) triiodothyronine, as a mitogen for keratinocytes; and (5) cholera toxin, to upregulate cAMP levels. This

1042-3699/02/$ – see front matter D 2002, Elsevier Science (USA). All rights reserved. PII: S 1 0 4 2 - 3 6 9 9 ( 0 2 ) 0 0 0 1 0 - 9

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method allowed keratinocytes to grow into a large epithelium sheet — 10,000 times larger than the original biopsy — in a short time period, enough to cover the large body surface areas. Investigators were then able to release sheets of keratinocytes by using the enzyme dispase, which could digest adhesive molecules holding the keratinocytes to the plastic substrate without digesting the cohesive links between the keratinocytes themselves. O’Connor et al. [39] reported the first human use of autologous cultured keratinocyte sheet grafts for burn wounds. Cultured keratinocyte sheets have been applied in various clinical scenarios such as chronic skin ulcers [51], congenital nevi [20], and junctional epidermolysis bullosa [5]. This approach revolutionized surgical treatment of burns and was the forerunner to the impetus to engineer skin to enhance tissue regeneration and repair. Two major approaches are currently used for in vivo tissue engineering of skin and mucosa. The older technique of Rheinwald and Green [47] is based on the use of a serum-containing medium and a feeder layer of lethally irradiated transformed cell line of mouse fibroblasts. The second approach relies on serum-free media in the absence of a feeder layer. The presence of irradiated feeder cells and serum can be a confounding factor in keratinoctye cultures when used for experimental research. This technique also would have difficulties receiving Food and Drug Administration (FDA) approval because of the potential transmission of unknown elements such as slow viruses to graft recipients. For this reason, many investigators have tried to avoid using feeder cells and to reduce the amounts of serum and additives, such as pituitary extract. The various methods for in vitro engineering of the epidermis require the use of (1) coating of culture surfaces with molecules found in the extracellular matrix, such as collagen, fibronectin, or laminin, that assist in simulating in vivo conditions and (2) variable concentrations of calcium in the culture medium. Calcium ions play a vital role in the growth and differentiation of keratinocytes. Increasing calcium concentration is accompanied by an increasing level of keratinocyte differentiation, as evidenced by increasing numbers of formed desmosomes and the formation of multilayers and sheets of cells. In vitro engineering of the epidermis also requires (3) the use of complex biologic extracts such as bovine pituitary extract and (4) the addition of mitogens or trace elements. For a skin or mucosa substitute to obtain FDA approval, it is necessary to eliminate the use of serum and xenogeneic feeder layers, undefined biologic

extracts such as pituitary extract, and products that may be modes of transmission of diseases, such as any bovine products that are not certified from diseasefree herds. Several skin substitutes are approved by the FDA and are commercially available. Epithelium Epicel (Genzyme Tissue Repair Corp, Cambridge, MA) is composed of cultured epithelial autogenous keratinocyte sheets and has been commercially available since 1988. The procedure of generating cultured epithelial sheets follows the technique of Rheinwald and Green [47]. The epithelial sheets are composed of several stratified layers of keratinocytes that are formed by culturing submerged cells (totally covered by medium) for 10 to 15 days. An epithelial sheet produced by their culture system has several advantages. The first advantage is the possibility of a large expansion from a small donor site of 2 cm2 up to 10,000-fold, which could cover a full-surface body area of an adult, approximately 2 m2. The second advantage is the low risk of transmission of viral diseases such as bovine spongiform encephalopathy. Early clinical studies of epithelial sheets used for the treatment of burn wounds had encouraging outcomes [19,37]. Later clinical experiences, in contrast, demonstrated low graft ‘‘take’’ rate [16,48]. Disadvantages of cultured sheets of grafted keratinocytes include (1) time-consuming growth of keratinocyte sheets in which, even under optimal conditions, it would take 2 to 3 weeks to obtain a sufficient amount of keratinocyte sheets for grafting; (2) the use of potentially immunogenic materials in culture, such as serum, various additives (pituitary extract), and a xenogeneic feeder layer that may contribute to graft loss; (3) widely reported variations in the ‘‘take’’ of keratinocyte sheet grafts, which is lower than that of split-thickness grafts; (4) graft instability, such as graft fragility and blister formation, that may be secondary to the absence of rete ridges and inadequate formation of epidermal keratinocytes’ anchoring fibrils; (5) wound contraction, which is secondary to a difference in specific characteristics of the keratinocytes within the cultured sheets to that of normal in situ keratinocytes and to the lack of dermal component within cultured epithelial sheet graft [10,44]; and (6) the cost of the grafts production. Dermis Dermis plays several biologic and functional roles in the skin. The most important roles are (1)

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providing mechanical support for cells involved in skin structure formation, immunity, nutrition, and sensation; (2) providing skin elasticity and tensile strength; and (3) functioning as an anchor for epithelial glands and keratinizing appendage structures of the skin (hair, nails). During skin healing, the presence of dermis also supports faster reepithelialization, inhibits wound contraction, and improves esthetic outcome. Increased understanding of dermal structure and composition has guided the development of artificial dermal substitutes. Structure is only one of the material property requirements. Dermis also should be pliable, hemocompatible, minimally immunogenic, and eventually degradable and must minimize fluid loss and reduce scarring and contractures. In designing a dermal replacement, one must take several clinical observations into consideration: (1) the thicker the dermal layer of a split-thickness skin graft, the less the graft contracts; (2) full-thickness skin grafts contract minimally; (3) full-thickness dermal injuries heal by contraction and hypertrophic scarring, which produce subepithelial scar tissue that is nothing like the original dermis; (4) partial-thickness wounds with superficial dermal loss heal with less hypertropic scarring; and (5) the length of illness in burn cases is essentially restricted to the length of time the burn wound is open. One might hypothesize from these observations that the dermis provides information to the wound that modulates the healing process. If so, then a dermal replacement should provide the information necessary to control the inflammatory and contractile processes and the information necessary to evoke ordered re-creation of autologous tissue in the form of a neodermis. The initial replacement material also should provide immediate physiologic wound closure and be eliminated once it has provided sufficient information for reconstitution of a neodermis. Yannas, Burke, and colleagues [4,22,55] focused on these observations and developed a bilayered twostage model that resulted in a FDA-approved product that is marketed as Integra (Integra Life Sciences Corp, Plainsboro, NJ). Integra has a top layer of a silicone elastomer and a bottom layer of a porous network of cross-linked collagen and glycosaminoglycan. The rationale for the top silicone elastomer is to control bacterial ingress and water evaporation and provide additional mechanical support. The bottom layer is designed to ensure rapid wound adherence. Optimal porosity of this dermal analogue allows for its slow biodegration and the induction of vascular and cellular ingrowth that eventually replaces the dermal matrix with neodermis. Several weeks later,

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when fibrovascular ingrowth of the dermal analogue has occurred, the epidermal silicone layer must be removed in the operating room and the wound closed with a thin sheet of autologous skin [4,22,55]. Clinical studies with this material have been successful [21,49] and were approved by the FDA in 1997. Dermagraft (Advanced Tissue Science Inc, La Jolla, CA) [15,38] is another dermal substitute that uses a resorbable matrix material that is similar to materials used in suture fabrication. This material is composed of a polylactin acid on which dermal fibroblasts are incorporated. Once incorporated within the matrix, the fibroblasts secrete growth factors and extracellular matrix proteins. A dermal allograft harvested from a living or cadaveric donor is another possible dermal substitute. Unlike the rejection of allogeneic keratinocytes on transplantation, allografts of dermis seem to be transplantable without significant rejection because of comparatively low immunogenicity of the dermal components. If rejection occurs, it is usually directed toward the passenger leukocytes and endothelial cells that line the blood vessels. The use of a de-epidermized and decellularized dermis further diminishes the allograft immunogenicity. Allogeneic, acellular dermis prepared this way retains the structural architecture of the remaining dermal matrix. This dermal matrix has been shown to support fibroblast ingrowth, neovascularization, and keratinoctye migration from an overlying split-thickness skin graft or from seeded cultured keratinocytes [29]. A product on the market, AlloDerm (LifeCell Corp, Branchburg, NJ), is an acellular, nonimmunogenic dermis that retains the extracellular matrix structure and an undamaged basement membrane complex. It also possesses an intact vascular channel network that allows ingrowth of fibroblasts and endothelial cells from the underlying tissue. Bilayers of epithelium and dermis In contrast to the materials science and engineering approach of Burke and Yannas et al. [4,22,55], Bell et al. [2] took the approach of reconstituting dermal wounds by applying a preformed tissue. They started with the observation that fibroblasts introduced into a collagen gel would proliferate and reorganize the collagen into a contracted matrix containing exogenous collagen and the collagen and matrix proteins produced by the introduced fibroblasts. The rate and final extent of contraction varied inversely with the protein concentration and directly with cell number introduced into the gel. The resulting product is described as a dermal equivalent, which, unlike

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Integra, relies on living cells in tissue culture to organize the collagen network. The exact fiber structure and its relationship to normal dermis are not known. Subsequent experiments demonstrated in animals that these collagen gels reorganized by fibroblasts could be grafted onto full-thickness injuries and that they would support the growth of keratinocytes into an epidermal equivalent. The dermal component is composed of type I bovine collagen that has been organized by introduction of human fibroblasts. Foreskin keratinocytes were seeded onto the surface of the dermal equivalent. After several days of submerged culturing of the skin equivalents, cultures are then air-exposed to allow the epidermis to stratify, differentiate, and form a cornified layer [17,43]. The total manufacturing period is approximately 20 days. To date, clinical evaluation of this type of skin equivalent has not been reported in burn patients, although several in vivo animal studies have been conducted [35]. A bilayered human skin equivalent, Apligraf (Novartis Pharmaceuticals Corp, East Hanover, NJ), already has been approved by the FDA for venous ulcers and is likely to be commercially available for burn wounds.

Oral mucosa substitutes Preprosthetic and reconstructive oral and maxillofacial surgical procedures often produce open wounds in the oral mucosa. These wounds should be covered by a graft to prevent microbial infection, excessive fluid loss, foreign material contamination, or relapse (secondary to wound contracture) and assist in the prosthetic reconstruction of the patient and in the promotion of wound healing [13]. Currently, oral mucosa or skin grafts are used for this purpose; however, both of these grafts require a second surgical procedure and have disadvantages in intraoral use [34]. Oral mucosa is an excellent intraoral graft material but is available in a limited supply [31,34]. Split-thickness skin grafts are available in ample supply but may contain adnexal structures, and they express a different pattern of surface keratinization that can lead to the development of abnormal tissue texture in the oral cavity that could interfere with function [12,34,36]. The elective nature of most oral and maxillofacial surgical procedures should allow the flexibility and timing to develop an ex vivo tissue engineered oral mucosa that could be used for intraoral grafting procedures. The recent developments of oral keratinocyte culture techniques have paralleled those of skin keratinocytes [24], which has enabled the development of tissue-engineered autogenous oral

mucosa that is suitable for intraoral reconstructive procedures [25]. Structural and functional differences between skin and oral mucosa The wet environment of the oral cavity complicates reconstruction with skin grafts. The keratinized surface of grafted skin tends to macerate and become easily infected. Oral mucosa is different from skin in that it has a moist surface and lacks adnexal structures such as hair and glandular elements. Grafting of skin into the oral cavity can be complicated by the presence of adnexal structures, which can be seen as hair growth within the mouth. Oral mucosa, unlike skin, presents three structural variations that are located in specific anatomic locations within the mouth. These layers are (1) masticatory mucosa (ortho or parakeratinized; hard palate, attached gingiva), (2) lining or alveolar mucosa (nonkeratinized; lip, floor of mouth, cheek), and (3) specialized mucosa (taste buds; dorsal surface of tongue). The keratinized and nonkeratinized mucosa differ in the composition of their cell layers. In keratinized mucosa, the suprabasal cell layer is divided into three layers and designated spinous cells, granular cells, and keratinized layers with the major cytoskeleton keratin of 1/10 [28]. Typical keratinized mucosa possesses ‘‘keratohyalin granules’’ in the granular cell layer. Tonofibrils, aggregates of keratin filaments, frequently are seen in the cellular cytoplasm. In contrast, the suprabasal layer of nonkeratinized mucosa is less evident and ordered than that seen in keratinized mucosa. The layers are designated as spinous cell, intermediate cell, and surface cell layer, in which the major intermediate filament of keratin is 4/13. Neither keratohyalin granules nor prominent aggregates of keratin are seen in nonkeratinized mucosa. In vitro culturing techniques Most investigators and maxillofacial surgeons have used an irradiated layer of a transformed 3T3 fibroblastic cell line as a feeder layer to propagate and expand their oral keratinocyte population to generate oral keratinocyte (epithelial) sheets for intraoral grafting [11,45,52,54]. The oral mucosa epithelial sheet grafts were placed onto the periosteum of the labial aspect of the anterior mandible to assist in performing a vestibuloplasty. All of the studies demonstrated successful clinical outcomes and histologic findings of postgrafting biopsies, the longest of which was 4 months postoperatively.

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Normal epithelial layer was regenerated on the graft sites. Hata et al. [21] and Ueda et al. [53] reported that oral mucosa keratinocytes grew more rapidly and differentiated less than skin keratinocytes. Dermal substitutes for burn injuries also have been applied into the oral cavity, such as the bilayered membranes with a collagen-GAG/silastic sheet, similar to the Burke and Yannas’ Integra [4]. Although they showed successful postoperative appearances and an advantage of easy sterilization and cost effectiveness, in the authors’ clinical experience this material is difficult to handle. The silastic sheet does not present a problem with handling, but the collagen sponge becomes ‘‘sticky’’ when it absorbs blood, which results in difficulty during suturing. The environment of the oral cavity, a moist area laden with bacteria and lytic enzymes, may not be conducive to the collagen-rich dermal components used in skin equivalents. An oral mucosa equivalent not only must be anatomically similar to mucosa but also must possess the mechanical and handling characteristics of the mucosa to be useful within the intricate confines of the oral cavity. There have been reports of ‘‘oral mucosa’’ equivalent-like Apligraf [40,41]. So far, these equivalents are still experimental and have not been used in clinical studies. Another type of ‘‘oral mucosa’’ equivalent composed of de-epidermized dermis and cultured oral mucosa keratinocytes from buccal mucosa and hard palate was studied in Korea [6,7]. These oral mucosal substitutes were developed for toxicologic and pharmacologic studies and not for use in a clinical setting. Studies have shown that the concurrent grafting of a dermal component aids in enhancing the quality and time of wound healing [26,32,33]. Parenteau et al. [43] showed that the rate of closure of the wound and the increase in the percentage of wound repair is enhanced with the presence of dermis. The maturation process and biologic events of skin regeneration also are accelerated with the presence of a dermal substrate [9]. Inokuchi et al. [23] have found that autogenous fibroblasts within the grafted dermal matrix facilitated the longterm maintenance of the reorganized cultured epidermis by supporting self-renewal of the epithelium in vivo. Clugston et al. [8] noted that the absence of a grafted dermis resulted in a contracture of cultured keratinocyte autografts on the order of 50%. The development and grafting of a full-thickness oral mucosal graft with a dermis can assist in epithelial graft adherence, minimize wound contraction, and assist in epithelial maturation while encouraging the formation of a basement membrane [18]. Auger et al. [1] showed that a dermal equivalent would be

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best made out of human, rather than animal, collagen. The human collagen (dermis) helps to promote deposition of additional basement membrane constituents, which results in a more optimal pattern of keratinocyte differentiation and less immunogenicity than animal collagen. The authors have been successful in their own laboratory in the ex vivo production of an oral mucosal equivalent (EVPOME) using oral keratinocytes seeded onto a human cadaver dermal matrix, AlloDerm [24,25]. AlloDerm is an acellular, biocompatible, human connective tissue matrix with an unaltered extracellular matrix and intact basement membrane, which consistently integrates into the host tissue. Most importantly, AlloDerm trims, adapts, and sutures like autologous tissue. Human de-epidermized dermis that has retained its basal lamina, consisting of keratinocytes combined with a mesenchymal or dermal component, has successfully shown enhanced epithelial morphogenesis and an increase in expression of differentiation markers when it is grown at an air-liquid interface [46]. Tissue-engineered oral mucosa Most reconstructive procedures in oral and maxillofacial surgery are of an elective nature. This gives surgeons the ability to time the biopsy of autogenous mucosa with the need of a sufficient size of tissue necessary for the planned surgical reconstruction. In developing a methodology to engineer any tissue, it is necessary to abide by the requirements and restrictions imposed by the FDA. The cultivating technique of Rheinwald and Green [47] uses a xenogeneic irradiated fibroblast cell line, 3T3, as a feeder layer to enhance keratinocyte growth. During the culturing period to expand human cells they are exposed to a transformed murine cell line. This contact potentially could contribute to cross-examination or transfection of the mutational or xenogeneic DNA into the cocultured human keratinocytes. Serum and a xenogeneic feeder layer contain undefined factors such as slow viruses (Creutzfeld-Jakob disease, ‘‘mad cow’’ disease, or bovine spongiform encephalopathy) and foreign contaminants [50]. The importance of not using a feeder layer and serum to culture oral mucosal autografts is obvious, especially in elective surgery because of the potential of the introduction of undetermined risks to the patient. Other investigators also support this point [14,27,43]. In our approach to tissue engineering an oral mucosal equivalent we use a serum-free culture system without a feeder layer. We also have successfully eliminated the use of bovine pituitary extract in the medium, thus having a completely defined

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Fig. 2. Transmission electron micrograph of keratinocytes in D4E. Numerous desmosomes (arrows) are formed between keratinocytes, while abundant tonofibrils are seen in the cytoplasm of the keratinocytes (osmium tetroxide postfixation and uranyl acetate/lead citrate, original magnification 17.000).

culture medium for the manufacture of their ex vivo produced oral mucosal equivalent (EVPOME) [24,25]. The authors’ EVPOME is composed of autogenous oral keratinocytes and a cadaver acellular, AlloDerm (Fig. 1 A – C). Electron microscopic evaluation of the EVPOME shows that the AlloDerm retains an intact basement membrane and anchoring fibrils on the papillary surface [29]. After being cultured 4 days submerged, the authors’ EVPOME shows several layers of keratinocytes adherent to one another via desmosomal attachments (Fig. 2), whereas specific junctional structures between basal cells and the basement membrane of the AlloDerm were not seen at that time (Fig. 3). At day 11 EVPOMEs, cultured 4 days submerged and 7 days at an air-liquid interface, numerous rudimentary hemidesmosome-like structures were seen incorporated into anchoring fibrils of the basement membrane of the AlloDerm (Fig. 4). This finding seems to indicate that the basal cell layer was attached firmly to the underlying dermal equivalent of the day 11 EVPOME, suggesting an ability of the epithelial layer to withstanding shear stress.

From a 4  4 mm2 punch biopsy of the palate it would take approximately 40 days to fabricate an EVPOME the size of one US dollar bill. This size EVPOME should be large enough to cover most mucosal defects. Approved human clinical trials were initiated in the Fall of 2000 at the Dental School Hospital of Niigata University, Niigata City, Japan. Our group at the University of Michigan also is in the process of obtaining FDA approval for a tissue-engineered oral mucosa for use in human clinical trials. The clinical protocol that was used for the first patients in the study performed at Niigata University in Japan was first to take a 5  5 mm2 punch biopsy of the retromolar trigonal mucosa in an outpatient setting under local anesthesia. The biopsy is planned sufficiently before the surgical procedure to ensure that an adequate piece of EVPOME is available for grafting. In most cases, to date, a period of 4 weeks has been sufficient. Oral keratinocytes are dissociated from the biopsy and expanded in a standard, serumfree defined culture medium. Once a sufficient number of oral keratinocytes has been harvested, 1.25 

Fig. 1. (A) Ex vivo produced oral mucosa equivalent (EVPOME) cultured 4 days submerged (D4E). Continuous epithelial monolayer has developed over dermal component, AlloDerm (Life Cell Corporation, Branchberg, NJ; H&E, original magnification l25). (B) EVPOME cultured 4 days submerged and 7 days at an air-liquid interface (D11E). Epithelial layer of D4E has started to stratify and differentiate. Keratinocytes in superficial layer are flattened and eosinophilic (H&E staining, original magnification 250). (C) EVPOME cultured 4 days submerged and 14 days at an air-liquid interface (D18E). An increase in stratification of the layers is noted that is consistent with a more fully differentiated epithelium. Epithelial layer demonstrates parakeratinization (H&E, original magnification 150).

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Fig. 3. Transmission electron micrograph of dermal-epithelial junction in D4E. There are no specific junctional apparatus seen between the basal cells and basement membrane. Original, retained, anchoring fibrils (arrowheads) in the papillary surface of AlloDerm (osmium tetroxide postfixation and uranyl acetate/lead citrate, original magnification 17.000).

105 cells/cm2 are seeded onto the acellular cadaver dermal equivalent, AlloDerm. The protocol outlined by Izumi et al. [25] is then followed. Briefly, the composites of oral keratinocytes and AlloDerm are cultured submerged for 4 days and at an air-liquid interface for 7 days to encourage epithelial stratification (Fig. 1 B). This protocol was determined to be optimal through in vivo grafting studies performed in SCID mice (Izumi et al., Tissue engineering, 2002, manuscript accepted for publication).

In patients, the EVPOME is produced and transplanted on day 11 after the oral keratinocytes have been seeded onto the AlloDerm. A gauze bolster or stent is then used to stabilize the EVPOMEs at the time of surgery. Surgical stents or bolsters are removed at 6 days postoperatively, and the surface of the transplanted EVPOME at the time is scraped with a swab for cytologic examination. The presence of small, round-shape cells suggests the presence of basal cell-like characteristics. Transnasal feeding is

Fig. 4. Transmission electron micrograph of dermal-epithelial junction in D11E. Hemidesmosomal-like structures (arrows) incorporated into anchoring fibrils are well developed. Note anchoring fibrils (arrowheads) newly integrated within the hemidesomosomal-like structures (osmium tetroxide postfixation and uranyl acetate/lead citrate, original magnification 30.000).

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used until removal of the stitches to minimize disruption of the grafted EVPOME. A soft diet is begun at postoperative day 8 at the time of removal of the transnasal feeding tube. At 4 weeks postoperatively, a punch biopsy is performed for histologic examination. In some cases, AlloDerm without autogenous keratinocytes as a control also has been transplanted onto oral mucosal defects. In contrast to the EVPOME, the AlloDerm graft without an epithelium showed more shrinkage over time postoperatively, which resulted in a greater degree of wound contraction. The AlloDerm graft without an epithelium caused an indurated wound, which could impair soft tissue mobility. On histopathologic examination of 4 weeks after surgery, the epithelial layers of the EVPOME and AlloDerm without epithelium demonstrated a regenerative, well-stratified epithelial layer. The presence of endothelial cells was evident as was a marked vascular ingrowth and cellular infiltration into the underlying dermal component of the EVPOME and AlloDerm alone. Because the presence of the formation of intense granulation tissue may lead to additional scarring, the grafted AlloDerm without an epithelium might result in a functional compromise within the oral cavity. The histopathologic features of grafted EVPOMEs showed a favorable remodeling and incorporation within the host tissue during the healing phase. Studies are in progress using tissue-engineered oral mucosa as a vehicle for the use of gene therapy to enhance wound healing and/or transmucosally administer systemically needed growth factors.

Summary To date, successfully developed EVPOMEs in a serum-free culture system without a feeder layer are the most acceptable and promising oral mucosal substitutes for human intraoral grafting because of minimal risk of foreign contaminants, easy handling and stitching, subsequent rapid revascularization into dermal component after transplantation, and contribution to favorable open wound closure without functional compromise, although several types of oral mucosal substitutes described in this article have been used in patients.

Acknowledgment The authors thank Masaaki Hoshino for his technical assistance, Dr. Michiko Yoshizawa for her input and involvement in the development of our

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tissue-engineered human oral mucosa, and Dr. Cynthia Marcelo for many fruitful discussions.

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Materials and techniques in maxillofacial prosthodontic rehabilitation Heidi Huber, DMDa,*, Stephan P. Studer, DMD, PhDb a

Regional Center for Maxillofacial Prosthetic Rehabilitation, University of Pittsburgh Medical Center, 2-South Montefiore Hospital, 3459 Fifth Avenue, Pittsburgh, PA 15261, USA b Department of Cranio-Maxillofacial Surgery, University Hospital of Zurich, Zurich, Switzerland

This article summarizes the English literature about current materials and techniques applied in maxillofacial prosthodontic rehabilitation that were cited in the MEDLINE database between the years 1990 and 2000 (Table 1). The evaluated articles were categorized into three different groups: (1) reviews, case reports, editorials, (2) laboratory studies and animal studies, and (3) human trials, either with or without control group or randomization. Case reports, review articles and editorials clearly predominated the literature search with two thirds of the publications, followed by laboratory and animal studies with one fifth of the articles. Only 14%, or every seventh of the published articles about maxillofacial prosthodontics, reported results in the form of human trials. The presented uneven distribution in favor of anecdotal reports rather than scientifically based hard data reflects the typical character of modern maxillofacial prosthodontics. It is definitely the art and science to rehabilitate esthetics and function of patients with acquired, congenital, and developmental defects of the head and neck [4]. Similar findings with a predominance of case reports and a small number of human trials were reported elsewhere, recently [141]. This article concentrates on the scientific facets of maxillofacial prosthodontic rehabilitation based on laboratory, animal, and human studies. Developments in materials and techniques that occurred during the last 10 years were subdivided into (1) materials and

* Corresponding author. E-mail address: [email protected] (H. Huber).

techniques with conventional facial prostheses and elastomers, (2) recent developments with acrylic resins, (3) presurgical orthopedics in cleft palate patients, (4) conventional obturator prostheses (surgical, interim, definitive), (5) conventional palatal lift prostheses, palatal augmentation prostheses, and speech aid prostheses, including prosthetics with glossectomy and mandibulectomy patients, (6) implants in maxillofacial prosthodontic rehabilitation, and (7) financing, costs, and quality of life.

Materials and techniques with conventional facial prostheses and elastomers Maxillofacial prosthodontic rehabilitation and materials Patients present with facial defects that are the result of congenital anomalies, trauma, resection of cancer, or some combination of these occurences. Surgical reconstructive methods are constrained by the integrity and availability of local and remote tissue beds, the patient’s ability to withstand such procedures, the difference in tissue colors and textures that become obvious after surgical procedures, and the requirements of viewing a site that has undergone surgery because of cancer. Prosthodontic reconstruction is constrained by the patient’s ability to accept something that is not natural and is removable, the resulting tissue bed that holds the prosthesis (i.e., mobile tissue that emphasizes lines of demarcation of a prosthesis), the retention of a prosthesis, and the materials currently available.

1042-3699/02/$ – see front matter D 2002, Elsevier Science (USA). All rights reserved. PII: S 1 0 4 2 - 3 6 9 9 ( 0 2 ) 0 0 0 1 8 - 3

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Table 1 Distribution of the selected English literature about maxillofacial prosthetics between the years 1990 and 2000 with its predominance of case reports and review articles in comparison to the small number of human trials Distribution of publications 1990 – 2000 Implants in MFP Conventional obturator Palatal lift, augmentation and speech aid prostheses; glossectomy patients and mandibulectomy patients Conventional facial prostheses Elastomers, polymers, silicone Cleft palate and MFP Acrylic resin in MFP Education, cooperation History, general reviews about MFP Financing, reimbursement Total publications

Case report, review editorial

Laboratory animal study

Human trial, noncontrolled

Human trial, controlled

Total

29 56 10

3 5 0

4 7 3

4 4 0

40 72 13

21 4 X X 8 12 0 140

1 33 X X 0 0 0 42

1 1 X X 0 0 2 18

0 1 X X 0 0 2 11

23 39 X X 8 12 4 211

Currently, osseointegrated implants have improved retention significantly and, in some cases, appearance and acceptance of prosthetic reconstruction, when a choice is available to the patient. Surgical versus prosthetic rehabilitation Typically, prosthetic reconstruction is indicated with larger defects if they involve cancer and some sort of adjunctive therapy that require some time to evaluate for recurrence. Decreased vascularity and fibrotic tissue at the defect periphery that result from radiation treatment may preclude surgical rehabilitation completely. In situations that involve intraoral and extraoral tissues or selected facial structures (i.e., ears), surgical procedures may not provide predictable functional or esthetic results. Age is another factor involved with the choice of prosthetic versus surgical reconstruction; older patients are more likely to receive restoration via prosthetic means. With respect to facial defects, DaBreo et al. [30] discuss surgical and prosthetic considerations in the management of orbital tumors that can apply to other facial defects. They mention, however, that the deliberation of saving the globe when there is malignant penetration of the bone but not the periorbita should be based on the following factors: (1) tumor histology, (2) contralateral vision, (3) the probability of permanent diplopia, and (4) the need for irradiation in doses that would destroy vision. The complications involved with irradiation of orbital tumors present surgical resection as the desired treatment most of the time. This choice ultimately gives the patient the option of prosthetic rehabilitation, the success of which depends on the size of the defect,

the tissue health and integrity, and the method of retention [30]. Although size of the defect is always at issue, often prosthetic rehabilitation is the option of choice for larger defects. Dacron polyurethane has been used as a custom-made prosthetic appliance with autologous bone in large frontoorbital defects, however. Autogenous bone or alloplastic implants by themselves may not present the best result contour symmetry, strength, or resistance to infection; however, the combined efforts of both materials have been in use since 1986 [111]. Maxillofacial prosthetic materials Currently the primary prosthetic materials available and under investigation are metals and polymers independently or in combination. Dispensing with the issues of prostheses not being from ‘‘Mother Nature’’ and the subsequent acceptance or lack thereof by each recipient, a review of some of the physical and biologic properties of these materials and how to improve upon them with respect to their clinical use as maxillofacial prostheses follows. Elastomers Silicone The four categories of silicone are based on application: 1. Implant grade: This material must at least meet or exceed Food and Drug Administration requirements. The materials used in breast implantation have led to investigation

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2.

3. 4.

(medical and legal) into the cytotoxicity of this material. More stringent testing is performed on this grade of silicone. Medical grade: This material is used externally and is found primarily in facial prostheses. Some studies test the cytotoxicity of this material; however, none has reported any negative side effects. Clean grade: This material applies to use with food coverage and packaging. Industrial grade: This material is mostly used for industrial purposes.

Silicone (or polydimethylsiloxane) is a name that is applied broadly to the primary material used in the fabrication of maxillofacial prostheses in which flexibility because of tissue mobility or undercuts is concerned or in which reshaping of the anatomy is needed prosthetically. Silicone and methyl chloride react to form dimethyldichlorosiloxane. When water is added to dimethyldichlorosiloxane, a fluid polymer, polydimethyl siloxane, is formed that is white and translucent and of varying viscosity, which is determined by the length of the polymer. Research on the modification of these polymer chains will lead to potential improvements in this material for maxillofacial prosthetic use. Cross-linking the long polymer chains at various points forms a network that is difficult to separate mechanically and makes it less likely to degrade from aging and various environmental factors. Room temperature vulcanizing silicones The silicones are also categorized into two types according to this cross-linking process, which is also known as vulcanization. Room temperature vulcanizing (RTV) silicones typically require a stannous octoate catalyst and an orthoalkyl silicate cross-linking agent (as in Silastic 382 and 399), in which polymerization occurs via a condensation reaction, or a chloroplatinic acid catalyst and hydro-methylsiloxane as a cross-linking agent (as in MDX 4-4210), in which the polymerization reaction is an addition reaction. Fillers such as silica or diatomaceous earth are often added to increase the tensile strength; however, a significant loss of translucency occurs with these fillers. This problem exists primarily for the Silastic 382, 399 products. The medical grade silicone MDX 4-4210 has been formulated with few fillers and has significantly more acceptable translucency without the loss of tensile strength. One of the sacrifices with MDX 4-4210 is the increase in time to cure at ‘‘room temperature’’ that may be decreased by bring-

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ing the temperature to 150°C for approximately 5 minutes [33]. Metal molds would be necessary in this situation, as opposed to stone with ambient temperature processing. In a survey conducted by Andres [7] regarding the use of extraoral maxillofacial materials, MDX 4-4210 has been used by 41% of clinicians surveyed. This material’s popularity is because of its improved physical properties, which have been investigated thoroughly [1,13,37,58,59,66,69,76,102,105,118, 134,139,140], such as increased resistance to tear, surface texture and Shore A hardness measurements being within the range of human skin and its compatibility with most skin adhesives [89]. It also has proved to be quite color stable [28]. Recent studies Some of the more recent studies on physical properties of the various silicones for maxillofacial prosthetic use have revealed that the physical properties change with the incorporation of coloring agents. In particular, Haug et al. [56] have reported that dry earth pigments, kaolin, and rayon flocking acted as a solid filler without bonding to the silicone, and artist’s oils and liquid cosmetics acted as a liquid base without bonding to the silicone matrix. Haug et al. [57] also proposed in another study that the addition of colorants could have a stabilizing effect on the elastomer color when it is exposed to weathering. Hulterstrom et al. [61] evaluated the changes in appearance of silicone elastomers for maxillofacial prostheses as a result of aging and found that the condensation-type polymers had an increased opacity in an aqueous environment, whereas the additiontype polymers showed the smallest color changes. Despite the higher content of filler in the additiontype polymers, they generally displayed lower opacity than the condensation-type polymers. Because of the higher viscosity in the condensation-type polymers, however, they may present a greater potential for intrinsic coloring. Mohite et al. [88] studied the effect of environmental factors on the physical properties of elastomers for facial prostheses, particularly two silicone elastomers (MDX 4-4210 and Cosmesil) and a polyurethane (Epithane-e). The test results indicated a significant difference in the tear characteristics between silicones and polyurethane and between the control and specimens exposed to environmental factors. These environmental factors affected the urethane the most and MDX 4-4210 the least. Chlorine and nitrogen dioxide exposure influenced degradation of Cosmesil and Epithane-3 to the point at which sample testing was impossible, and

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ultraviolet radiation exposure influenced degradation of Epithane-3 to the point that it could not be tested. Simulated sebum and ozone did affect Cosmesil and Epithane-3 with no trends observed. Wettability Another preferred physical property of any material that contacts tissue is ‘‘wettability.’’ This property is often measured by evaluating the contact angle and surface energies. The surface energy of a material informs us of the available energy for adhesion with mucosa or bacterial contact. Materials with high wettability establish a lubricating layer between the tissues that they contact and potentially decrease tissue discomfort caused by friction. Waters et al. [129] compared the wettability of various silicones with denture acrylic resin material and found no significant differences in the surface energies of the silicones, but all were significantly lower than denture acrylic resin. In another investigation of properties of maxillofacial silicone elastomers, Andreopoulos et al. [6] found that wetting properties degraded with increasing silica volume fraction. Evaluation of this property is necessary with each polymer modification in an effort to fabricate improved maxillofacial prosthetic materials. Newer silicone polymers (A-2186), Cosmesil, siphenylenes, and silicone copolymers The modification of the silicone polymer has led to some of the newer materials currently in use, such as A-2186 (Factor II) and Cosmesil. Factor II has displayed physical properties that were at least as acceptable or better than those of MDX 4-4210 [134]. These enhanced physical properties did not last after environmental factors were imposed, however, and MDX 4-4210 had superior physical properties relative to Factor II in a study conducted by Haug et al. [59]. Woolfaardt [135] discussed the advantage of Cosmesil’s ability to process to variable levels of hardness and have higher tear strength than MDX 4-4210. Veres et al. [125 – 127] also studied other properties of this material. Siphenylene polymer is another siloxane copolymer that has phenyl and methyl groups and exhibits much in the way of improvement with respect to improved edge strength, superior coloration, and low modulus of elasticity relative to other conventional RTV silicones in use for maxillofacial prosthetics [76]. Finally, some of the new avenues of current study involve silicone block polymers in which blocks of

polymers other than siloxane are positioned with the traditional siloxane polymers in an attempt to modify the current physical properties of conventional silicone. An example of this is the intertwining of poly methyl methacrylate into the chains of siloxane. This material is currently under study [117]. Silicone foam (Silastic 386) Silastic 386 was investigated by Firtell et al. [40] in 1976 for situations in which the weight of a prosthesis is at issue. Because of the size of some defects, the weight of the consequent prosthesis does not permit its wear. Firtell et al. tested the feasibility of combining foam RTV silicone rubber with conventional RTV silicone rubber to obtain a material that is lighter in weight. An element in silicone, when mixed with a stannous octate catalyst, releases a gas in the vulcanization process as bubbles are released with the resulting silicone mass being increased and density being decreased, which presents a much lighter material. This process requires special flasks to deal with expansion problems while the gas is forming during processing. The mold also requires venting for gas release and reduction of expansion of the prosthesis. Firtell et al. varied the mixtures of foam and conventional silicone rubber and established that molding, accuracy, texture, and color problems could be overcome reasonably. One way of overcoming the tear strength is to coat the foam with another silicone, which in the process creates a stiffer product. Reduction of weight proportionate to reduction of tear strength of the material formulated was not predictable, however. This material is not one of the most popular among maxillofacial prosthodontists [7], especially for facial prostheses. Application of the foam is reported by Casey et al. [21] for the fabrication of an intracavity radiation prosthesis and tissue compensator, and it is important to be aware of this product as a potential maxillofacial prosthetic tool. High temperature vulcanizing (HTV) silicones These silicones, in general, possess greater physical properties (especially tear resistance) than RTV silicones, with the disadvantages involving the ultimate esthetics of the resulting prosthesis and the processing requirements. The process of vulcanization requires greater milling of the solid HTV stock elastomer for mixing the catalyst for cross-linking and pigmenting [80]. This material may present as a one- or two-component system with a putty-like consistency in which the two primary catalysts in-

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volved are platinum salt (for addition reaction polymerization) and dichlorobenzoyl peroxide (for condensation reaction polymerization). Chalian et al. [23] discussed the milling of HTV silicone using a two-roll mill with a motor drive so that greater translucency can be obtained with increased dispersion of internal pigments used. The next processing step involves curing in a heat-transferring metal mold at high temperatures. The milling process significantly reduces the potential for trapping air and consequent porosity that frequently occurs with hand incorporation of pigments with RTV silicones. This material is not as elastic as MDX 4-4210 or other RTV silicones and is not as applicable in situations of mobile tissue beds. Dow Corning [33] proposes the addition of one of their products, a polydimethylsiloxane oligomer-Electronic Fluid 200, as a softening agent. This cross-linking with proportional series of silicone fluid oligomers was part of a comprehensive clinical application program initiated by the Veterans Administration to attempt to produce an improved material based on ideal physical properties, established by Rahn et al. [101]. Lontz et al. [78] reported key physical properties under the material name of polydimethyl siloxane to initiate a new product line for prostheses only in a series of publications in the VA Bulletin for Prosthetics Research [80]. Polyurethanes This material is chemically composed of an extended segment of diisocyanate groups and a segment of polyol groups (a mixture of polyesters) and an organotin catalyst for the polymerization process to occur [76,80]. As these segments are varied in proportion to each other, so varies the softness of the end-product for its intended application, because maxillofacial prosthetics tend to require greater softness and flexibility. Epithane-3 and Calthane are the only polyurethanes currently available for fabricating facial prostheses. These are three-component, room-temperature curing systems that carry with them a certain amount of technique sensitivities in their processing. In particular, the diisocyanate component is hazardous, toxic, and moisture sensitive. Water contamination results in bubbles, deficits, and questionability in the curing of the final prosthesis. In one of Gonzalez et al.’s first reviews of this material [51], the requirements of accuracy and care for successful processing are addressed. Gonzalez et al. [50] further studied the various physical and mechanical qualities of polyurethane with the parameters of surface hardness, modulus of elasticity, strength, percentage of elonga-

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tion, and strength-to-modulus of elasticity ratio. The results of the study confirmed that these physical and mechanical properties can be altered and customized to suit the prosthetic situation by varying the ratio of Part (A) to Part (B) and addition of catalysts. Compositions with low quantities of the isocyanate [Part (B)] and no catalyst reached or approximated those parameters proposed as ideal goals to stimulate living tissues. Goldberg et al. [48] discussed in greater detail the alipathic diisocyanate and polyether macroglycol polymerizations with various cross-link densities and OH/NCO (glycol polymer) ratios. Stoichiometries that yielded between 8600 and 12,900 g/mol/crosslink and an OH/NCO (glycol polymer) ration of 1:1 resulted in polymers with the low modulus yet high strength and elongation necessary for maxillofacial applications. This particular document also discussed the importance of proper mixing to avoid air entrapment and phase separation of the catalyst, which are conditions that may make the prosthesis less durable [50]. An et al. [5] established that there was no statistically significant difference on elastomer properties with respect to variations in curing temperatures and times. With respect to processing, molds of metal (if higher temperatures are used so that greater tensile strengths are obtained after polymerization), stone, urethane, and epoxy can be used. The moisture sensitivity of this material presents the need to dehydrate stone molds before use. Stone molds also must be fabricated from a wax as opposed to a clay sculpted contour to avoid contamination. Once the prosthesis is stained from such contaminants, the external staining does not adhere as well. The current pool of adhesive systems is not compatible with this material, and cleansing the adhesive frequently results in removal of extrinsic stain. Although the final prosthesis has the isocyanate in a bound and presumably nontoxic form, there is evidence that the ternary composition for maxillofacial prosthetics is toxic to human excised donor orofacial tissue cells [79]. Lontz et al. [78] have reported free isocyanates in cured restorations, which is an obvious concern regarding cytotoxicity and tissue irritation after prosthetic wear. Some of the positive qualities of this material pertain to its flexibility without losing strength at the edges, which allows margins to be made thin to obtain optimal esthetics. Goldberg et al. [49] evaluated the tear energy, or the energy necessary to generate a unit area of torn surfaces, for polyurethanes versus other commercial maxillofacial materials. The commercial materials had values between 0.63 and 6.56 erg/cm2, whereas the polyurethanes ranged from

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8.49 to 50 erg/cm2. The intrinsic and extrinsic coloring of this material is possible with esthetic results that exceed most currently used materials [22]. Chu et al. [25] investigated the effect that ultraviolet (UV) stabilizers had on the stability of mechanical properties of Calthane ND2300. The study indicated that the service life was prolonged; however, modulus of elasticity and tensile strength decreased after extended exposure to ultraviolet light. The rate of decrease was much less for specimens with UV stabilizers. The group that contained a UV stabilizer and an antioxidant was noted as the best of the groups tested with respect to the percentage of retention of tensile strength and modulus of elasticity. Elongation at break increased when the duration of UV aging was increased in all systems. The explanation for this phenomenon was provided in terms of the concepts of the fragmentation of macromolecules that resulted from UV aging.

necessitates the use of denture flasks. One factor to consider with this process on long-term patients is that duplication of the prosthesis must be handled at the mold fabrication level, because the mold is destroyed after the prosthesis is deflasked. This approach may be considered especially for patients with no major tissue changes anticipated near the defect, and periodic remakes may be part of the protocol of treatment. Light-cured acrylic The introduction of light-cured polymers has been a useful tool for provisional prostheses intraorally and extraorally. The characteristic brittleness, inability to stain, and poor thermal conductivity of these acrylics point to the use of methyl methacrylate as the acrylic of choice for most acrylic prostheses, however. Copolymer

Recent developments with acrylic resins Acrylic methyl methacrylate resin: heat-polymerizing and autopolymerizing Restoration of defects with tissue beds that are relatively nonmobile lends itself to the use of acrylic resin prostheses with nice esthetic results. Often acrylic is the material of choice for provisional prostheses, and modifications are readily made as tissue changes occur with sequential surgeries, healing periods, and radiation-related tissue changes. Modifications can be made by further addition or removal of acrylic, application of soft tissue conditioner as a temporary liner or functional impression material to be transformed into acrylic, and staining alterations. Margins can be made substantially thin with subsequent minimization of the appearance of demarcating lines. Compatiblility with most adhesives is good, and adjustments with respect to addition or removal of acrylic and stain are accomplished readily. Extrinsic and intrinsic staining can be accomplished by types of acrylic, although color stability in heat-polymerized acrylic is superior to that of autopolymerized acrylic when considering UV light exposure [16]. The tertiary amines that leech out of the autopolymerizing acrylic prosthesis are toxins that could be omitted by using a heat-processed acrylic. Processing Processing involves the use of creating molds out of dental stone that are placed under pressure, which

Polymer chemistry has the greatest potential of improving currently available materials that can be more clinically serviceable in maxillofacial prosthetics. In a conference on materials research in maxillofacial prosthetics in 1992 [8], new materials were being fabricated by which the polymerization processes were varied and more versatility was introduced into their clinical application. Combining acrylic polymers of high molecular weight with blocks of other polymers is an attempt made to obviate the problems encountered with the original acrylic copolymers. Palamed is one of the prototypes of plasticized methyl methacrylates, or copolymers, used in maxillofacial prosthodontics. The prostheses made of copolymers tend to be less durable and require more frequent repairs because of marginal tears, which are not successfully repaired. These restorations do not present with ease of fabrication, and with the esthetic results obtained, one would be served better by using another material. Over time the surface texture of these prostheses becomes sticky and tends to collect debris, which changes the appearance and ultimately creates problems with cleaning and retention of the prosthesis. Coloration Acceptance of a facial prosthesis ultimately is influenced by how realistically it is colored. The task of the prosthodontist is laid out in terms of matching tissue adjacent to a defect that is to be covered by

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a prosthesis. Skin is composed of many layers that vary in thickness and cellularity and pigments. Another factor to consider is surface texture, which also plays a role in light scattering. What the human eye sees when viewing human skin is a result of the light refracted and reflected by the different layers within the skin. Coloration is a highly subjective process that involves the perception of both the clinician and the patient and the type of lighting, which may vary from location to location. There are perceived gross color similarities among people within each racial group; however, complexities of skin patterns among individuals become apparent when we are involved with the task of color matching. The process of establishing a color for a prosthesis is essentially broken down into intrinsic and extrinsic staining of the prosthetic material. Intrinsic staining is the stage at which a base shade is established, whereas extrinsic staining involves further characterizing of the initial intrinsically stained prosthesis to attain esthetically optimal results. Workers have documented formulations of colorants in the formation of base shades and extrinsic staining for referencing purposes [2,12,19,34,41,47,82,99]. Beard et al.’s [12] classic article on coloration presents spectral measurements of the great variety in color of human skin. Attempts to standardize and quantify spectral values in terms of pigments used for various skin colors have helped to alleviate the prosthodontist of some, but not all, guesswork. The subjective biases also can be minimized by this process; however, this remains an artistic endeavor and always requires a skillful eye and hand for its ultimate esthethic acceptance. The literature presents numerous types of colorants for maxillofacial prosthodontic uses [11,17,19, 26,38,39,43,94,109,119]. Colorant selection depends on the provider and the material. Changes in physical and mechanical properties of certain materials, after addition of select colorants, have been reported. In 1984, Turner et al. [118] reported a statistically significant decrease in tensile and tear strength when kaolin and dry earth pigments were combined with polyurethane. Beumer et al. [16] reported that oil paints interfered with the setting reaction of MDX 4-4210. Intrinsic coloration The primary goal of intrinsic coloring is in imitating the physiologic colorings with pigments incorporated into a compatible commercial base material. Lontz et al. [77] discussed the physiologic details of skin in terms of colors: arterial red, venous redpurple, carotenoid yellow, melanoid brown, and opaque cellular lipids are colors to be replicated in a

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quantitative manner via a commercial Hunter digital color difference spectometer. These workers [80] present a range of intrinsic color shades to serve as a bulk coloration onto which extrinsic staining can characterize the prosthesis further. This type of technology is also useful in the commercial cosmetic industry, in which standard color references are required for different races and individuals. Hanson et al. [54] further discussed the application of commercial cosmetics with the coloring of facial prostheses. Some situations necessitate using cosmetics for the purpose of camouflaging lines of demarcation or color differences between a prosthesis and skin. This necessity is often influenced by seasonal changes and resulting skin color changes. Although coloration greatly influences prosthesis acceptability, translucency is also a major factor. Usually the clinician initially considers the inherent translucency of the base material. After the addition of selected colorants, variations in translucency can occur. Johnston, et al. [63] established a translucency parameter by collecting optical scattering and absorption coefficients and applying the Kubelka-Munk reflectance theory. The actual procedure involved measuring the color difference between a 0.13-cm thick sample of a colored medical-grade silicone sample, which was placed on an ideal dark backing and an ideal white backing. Significant differences were noted among the translucency parameters of the colorants. Another study that incorporated the Kubelka-Munk theory was performed by Ma et al. [83] for the evaluation of the Kubelka-Munk theory for agreement in color prediction of thick pigmented samples and linearity of optical absorption and scattering coefficients with concentration of colorant in a maxillofacial elastomer. The colorants tested were generic opacifiers, dry mineral earth pigments, and fibrous colorants. The agreement between theoretical and observed colors was significantly closer for the fibrous colorants than for the dry mineral earth pigments of the same labeled color. Color stability is an issue that has been investigated as part of the primary physical property study of base elastomers for maxillofacial prostheses. General degradation of the prosthesis may be related to multiple factors, with environmental weathering and aging being the key contributors. Lemon et al. [75] assessed the efficacy of an additive, intrinsic, broadspectrum UV light absorber on the color stability of pigmented elastomer. The material was weathered artificially and outdoors at exposure levels of radiant energy of 150 to 450 kJ/m2, with a perceptible color change reported. The artificial aging influenced a greater change than the outdoor aging, with the UV

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light absorber (UV-5411) not displaying any protection from color change. Tamamata et al. [115] study suggested that aging is the primary influence on color changes observed in HTV and RTV base polymers, as opposed to exposure to sunlight. Chu et al. [25] chose to investigate polyurethane and the efficacy of polymer additives, especially UV absorbers and antioxidants. Evaluation of 11 different types of UV absorbers was coupled with one antioxidant and incorporated into a polyurethane system that was exposed to an artificial UV source. The UV stabilizers enhanced the UV resistance; however, tackiness from UV aging resulted and could not be resolved satisfactorily. The problem of yellowing was improved significantly. The most promising UV absorbers were Tinuvin 770 and the combination of Tinuvin 328, zinc oxide, and an antioxidant. The use of ceramic pigments as applied to maxillofacial prostheses was evaluated by Seluk et al. [109]. Samples of unsintered versus sintered pigmented porcelain were mixed into Silastic 44210 and evaluated with respect to color stability after accelerated aging. There was clear demonstration that sintered pigments were more color stable than unsintered pigments. This is the only published article discussing this pigment with respect to color stability. The effects of pigmented porcelain on the physical properties of base materials have not been reported. Bryant et al. [18] attempted to evaluate the use of commercially available sunscreens with sun protective factor of 15 as a method of decreasing the color changes caused by UV radiation on an intrinsically colored elastomer (MDX 4-4210). The colorants talc and nylon flock were incorporated into silicone samples, which were divided into three groups based on the different types of sunscreens tested. The results indicated that none of the sunscreens provided any UV radiation protection to the silicone, and the addition of PABA caused a significant color degradation of the silicone. Extrinsic coloration The final esthetic characterization of any prosthesis, as in porcelain, acrylic, or elastomer, involves staining the external base material. This discussion is limited to elastomers involved with facial prostheses. Most extrinsic staining involves the use of a medical grade silicone adhesive that is thinned with xylene to a more liquid consistency. The pigments are then distributed into small amounts of this mixture on a palette for application to the prosthetic surface. Characterization also involves surface texturing of the base material that will be colored. Polymerization of the adhesive (typically 30 – 40 minutes) can be

hastened by heat because xylene is evaporated through this process. Silicone prostheses form a covalent bond with the silicone adhesive, which becomes part of the prosthesis. This process is termed ‘‘integral extrinsic covalent bonding,’’ which is modified by surface texturing [80]. Adjustments can be made by reapplication if more color is needed or conversely by removal of excess color with chloroform. This method of extrinsic coloration is relatively easy and effective. Detail textures in the base prosthesis are lost, translucency is diminished, and rubbing or peeling of color with bending of the prosthesis occasionally occurs. A modification of this procedure was introduced by Quellette [100] in which extrinsic coloring was sprayed onto the surface. This method does not permit good control of color application and only allows a general distribution of coloring. An extrinsic method of coloring that was introduced by Schaaf [107] is tatooing. A fine neddle penetrates paint approximately 2 mm below the prosthesis surface. After the various shades for characterization are prepared, they are painted on the surface of the prosthesis and incorporated into the base material with a tatooing machine. The excess paint is removed with xylene and soap and water. This method eliminates the color stability problem and maintains the base prosthesis surface texture without altering translucency. This is a techniquesensitive and time-consuming process that requires some learning, however. It is difficult to correct for excessive coloration, and because the prosthesis is punctured, there is a weakening effect, especially at thinner regions of the prosthesis. Retention Unless a prosthesis (introral or extraoral) is retained it cannot be successful. The literature displays several methods of retention: natural teeth, clasps, precision attachments, nonrigid attachments, magnets, anatomic undercuts, elastics, eyeglasses, intraoral prostheses, adhesives of various forms, endosseous implants, and titanium hollow screw reconstruction plate systems. MacEntee [84] presented the use of nonrigid attachments placed on select teeth for added stability and retention of a maxillary denture combined with a nasopharyngeal obturator. Frederick [44] described a technique for the fabrication of a sectional interim maxillary obturator with retention augmented by a magnet. Firtell et al. [42] measured the influence of a simulated obturator on the amount of force required to dislodge a simulated unilateral removable partial denture with various clasp designs. The presence of an obturator reduces

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the retentive capability of a removable partial denture. Lingual retention seemed to provide more resistance to displacement than buccal retention. Infrabulge clasp designs seemed to be more retentive than suprabulge clasp designs. King et al. [64] further expounded on the indications and rationale for using cast circumferential and wire clasps for obturator retention. This article justly recognizes that obturator retention is multifactorial and that an understanding and application of basic fundamental prosthodontic principles and clinical judgment are necessary. Parr [97] discussed the use of retaining a facial prosthesis by attaching the prosthesis to eyeglasses, which can then be attached to the ears. This ‘‘piggyback’’ method is also used more thoroughly with combination prostheses situations. Evans et al. [36] illustrated the prosthetic restoration of a large midfacial defect with the use of an intraoral obturator retained by implants and a large silicone facial prosthesis attached to the obturator. The range of retaining prostheses is wide. Facial prostheses and combination intraoral-extraoral prostheses are the two situations that require the most attention. Much effort is placed on the improvements of adhesives and the use of intraoral and extraoral implants. Parel [32] discussed the decreasing dependence on adhesives for retention of facial prostheses. Good communication with surgeons regarding essential tissue removal for the purpose of optimal esthetics and retention of prostheses has influenced this decrease in adhesive use. Despite the decreasing use of adhesives and the increasing incorporation of implants, which are discussed further in this article for maxillofacial applications, adhesives are still an important tool for most patients. Adhesives Skin adhesive is a temporary method of retention. Adhesives are expected to retain prostheses during ordinary and extreme facial expressions, build-up of sebaceous secretions and water, and change of weather conditions. Biocompatibility is another requirement that must be met because adhesives contact the skin for lengthy periods. Immediately traumatized tissue (via surgery or irradiation) must be treated with great care and the chemical composition of adhesives must be evaluated carefully before administration. Commonly used adhesives are categorized as: rubber-based liquid adhesives (natural and latex), pressure-sensitive bifaced tape, silicone, acrylic resin emulsions (gum mastics), and cyanoacrylates. This section is limited to the tape, silicone, and acrylic resin adhesives.

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Pressure-sensitive tape (double-coated polyethylene, 3M surgical tape) These materials are backing strips composed of cloth, paper, film, foil, or laminate coated with a pressure-sensitive adhesive. The adhesive is a rubbery type elastomer combined with a liquid or solid resin tackifier component, plasticizers, fillers, and antioxidants. Two advantages of tape are the ease of application and cleaning after removal. The bond is weaker than that of rubber adhesives. The primary indication for biphasic tape is with materials that have poor flexibility and nonmobile tissue beds. Silicone adhesives (Hollister) These adhesives are a form of RTV silicone dissolved in solvent. Once applied, the solvent evaporates and a tacky surface forms that can contact bond with another surface. Despite their low adhesive strength, they have good resistance to moisture and weathering with low water sorption. They are prone to dissolving in organic solvents such as xylene. Acrylic resin emulsions (Epithane-3, ProsAide) These adhesives are composed of acrylic resin dispersed in a water solvent that, when evaporated, leaves a rubber-like substance. Other materials within the mixture include synthetic rubber, vinyl acetate, reclaimed rubber, vinyl chloride, styrene, and methacrylic. Penetration and wetting can be controlled by addition of surfactants or altering the particle size of the dispersion. Increasing the viscosity can prevent penetration into porous surfaces. One surface must be permeable to water to dry the dispersion and develop the bond. In a survey [24] of 73 patients who wear facial prostheses, the ranking of preferences for various adhesive brands were (1) double-sided tape (41%), (2) rubber-based liquid (21%), (3) acrylic resin emulsions (19%), (4) silicone (4%) and (5) no reply. Some problems associated with adhesives are as follows: 

  



Patients with poor manual dexterity or coordination may not be able to apply the adhesive or position the prosthesis in a consistent manner. Margins adjacent to mobile tissue may require constant reattachment with facial movements. Allergic or irritational responses may persist. Poor hygiene may limit the wearing of a prosthesis because of interference with adhesive qualities. Some aromatic base adhesives may curl thin prosthesis margins.

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Routine removal of adhesive also may remove the external pigmentation.

The choice of a skin adhesive involves the status of the tissue and the material of the prosthesis that it contacts. Some adhesives bond more strongly with certain materials. Udagama [120] determined the best combinations to be polyvinylchloride with Epithane3 adhesive, acrylic with Medico, and polyurethane with Davol. Krill’s findings [71] are as follows: 

Silicone adhesive type B (aerosol) is the most effective for silicones.  Pressure-sensitive tape is the most effective adhesive for polyvinyl chloride.  Butyl alpha-cyanoacrylate is the most effective adhesive for polyurethane as tested for the Initiation of Peel test and Silicone Adhesive Type B for the Peel Effort test. Udagama [121] recommended the use of prefabricated urethane film (Factor II, Inc, Lakeside, AZ) a lining for prostheses. Because of the poor adheophilic properties of silicone, this urethane film permits the use of commonly used water-based skin adhesives such as Epithane-3 or 3M double-sided tape. These adhesives could be removed easily after soaking the prosthesis in water for approximately 10 minutes. A silicone prosthesis without a urethane lining first must be coated with a silicone-based adhesive followed by either Epithane-3 (Daro Products, Butler, WI) or Prosaide (ADM Tronics, North Vale, NJ). Controversy regarding the use of adhesives arises from the lack of information on the biomechanical performance of this material, and no standard exists for this material. Tam et al. [114] designed an apparatus for testing this material in which tensile, torsion, and combined tensile-torsion tests were performed. Wolfaardt et al. [135] further investigated the mechanical behavior of three adhesives through an in vivo pilot project on two subjects. Of the three adhesives tested—PSA1, Pros-Aide, and Dow Corning 355—Dow Corning 355 was statistically significantly stronger than the other two adhesives by means of tensile and combined tensile-torsion tests. Wolfaardt et al. cautioned the extrapolation of results of this study to clinical conclusions until more investigations are carried out. The testing of biphasic tape was performed by Polyzois [98], who measured the tensile bond strengths of five silicone facial elastomers to skin by use of five double-sided adhesive tapes. Significant differences were observed among the various silicone/tape combinations. Cosmesil and MDX

4-4210 elastomers had the strongest bond to skin with most adhesive tapes, whereas Silskin II, Cosmesil HC2, and RS 330T-RTV were the weakest. Tapes remained adhered to the skin and not to the silicone after bond failure. Haug et al. [60] investigated the percent adherent area and 10-minute and 8-hour peelstrength values as an in vivo evaluation of three adhesives with three prosthetic polymers. Hollister Colostomy Adhesive showed greater 10-minute peelstrength values when used with Silastic Medical Adhesive A. The peel-strength values of the Hollister Colostomy Adhesive when used with Silastic MDX 4-4210 increased with time. The greatest percent adherent was obtained with Silastic MDX 4-4210 using Pros-Aide Adhesive, and the lowest was obtained with Pros-Aide Adhesive using Silicone A2186 and Silastic Medical Adhesive A. There was no clearly superior adhesive/prosthetic material combination based on the test performed in this study. Individual skin variations must be considered, however. Primers were introduced as a material to aid in the improvement of bond between silicone and other prosthetic materials such as polyurethane liners. Wang et al. [128] evaluated two primers, three polymerization methods, and seven primer reaction times to determine the conditions for optimum adhesive bond strength. Bond strengths were significantly greater for polyurethane treated with primer 1205 rather than S-2260 regardless of the polymerization method or primer reaction time. No single polymerization method nor single primer reaction time consistently yielded greater bond strengths. Taft et al. [113] compared the adhesion-in-peel force of a silicone adhesive to autopolymerizing polymethyl methacrylate and light polymerizing urethane dimethacrylate gel with two surface textures: (1) pumice polished and (2) pumice polished and bead retention and two surface primers: (1) Dow Corning 1205 primer and (2) Dow Corning S-2260 primer. The 1205 primer produced an adhesion-in-peel force that was statistically significantly stronger regardless of the type of resin or surface preparation used.

Presurgical orthopedics in cleft palate patients Appliances Presurgical orthopedic appliances have been in use for at least 45 years [85] to facilitate the movement of alveolar cleft palate segments into a position that allows the plastic surgeon to repair the lip cleft more easily with an ultimate cosmetic improvement. The early orthodontic treatment of patients with

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alveolar clefts was demonstrated by Burston [20] as early as 1965 and was expanded later (1971). O’Donnell et al. [93] presented an analysis of the subject of presurgical orthopedics in the treatment of unilateral cleft lip and palate in 1974, and the use of these appliances has remained controversial ever since. A large international study of treatment outcomes in clefting had associated poor outcomes with active presurgical orthopedics [110]. A retrospective study in Toronto (using a sample of 40 adolescents with the repaired defect) presented that conservative presurgical orthopedics had no lasting effect on lip and nose esthetics and did not alter the need for subsequent revision surgery [103]. At the University of Pittsburgh, Winters et al. [132] in 1995 concluded that presurgical orthopedics does facilitate primary reconstructive surgery for the clefting and is considered essential in some of the new techniques of primary rhinoplasty. Various forms of presurgical orthopedic appliances, which fall into the three primary categories of (1) passive, (2) pin or screw-retained, and (3) active. All appliances require a maxillary impression to be made for the fabrication of cast and subsequent appliance. Passive appliance This appliance typically involves the use of a single acrylic base plate that passively covers the lateral alveolar cleft segments while the uncovered premaxillary segment is moved so as to be in as proper alignment as possible with the lateral segments through the use of external strapping. The patient is evaluated on a weekly basis for soft tissue irritation and guided movement of the alveolar segments, which is accomplished by selective additions and removal of acrylic from the appliance. Premaxillary movement is accomplished by varying the amount and direction of forces placed on the external strapping. This procedure depends greatly on the compliance of the caretakers of the infant, and the retention of the appliance is often poor. Pin-retained appliance A pin-retained appliance was introduced by Hagerty [53] in the way of an expansion bar of stainless steel (only), in which the pointed ends were inserted into the palatal bone for increased retention. Hagerty later added an acrylic base to the expansion bar. The method that is still in use currently was introduced by Georgiade et al. [45] in 1970, with a staple pin made from stainless steel wire for retention of acrylic plates that would cover the palate. In 1976, Latham et al.

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[74] presented an extraorally activated expansion appliance for cleft infants. This appliance contains an extraoral control knob that facilitates activation of the expansion bar, which leaves the anterior cleft regions free for premaxillary movement and clinical evaluation. Rapid expansion can be accomplished easily. Retraction of the premaxilla in a bilateral cleft lip and palate is accomplished by external strapping or an orthodonic elastic chain attached to a pin that penetrates the vomer. Varying the forces on the elastic chain guides the premaxillary movement. Latham et al. reported that it takes approximately 1 day for the vomer to respond to the force and pull into the midline. Once inserted, the expansion screw is then turned one-half revolution per day (or on alternative days), which can be accomplished by the parents. Each revolution provides 1 mm of expansion at the anterior ends of the expansion arms. Activation of the elastic chains must be performed by the plastic surgeon, who initially places the appliance. One of the concerns with this appliance is that the premaxillary pin can be pulled out. Alignment of the alveolar segments takes approximately 3 to 5 weeks, according to Latham et al.. When the alveolar segments are properly positioned, the surgeon removes the appliance and repairs the cleft lip either immediately or 2 to 3 days later. Active appliance A new procedure for nasal and alveolar molding has been introduced by the workers at the Institute of Reconstructive Plastic Surgery at New York University Medical Center [52]. This procedure involves reducing the size of the intraoral alveolar cleft and active molding and positioning of the nares tissues. A palatal appliance that covers the alveolar segments completely is worn by the infant until the alveolar segments are in a more favorable position. As material is selectively added to and reduced from the intaglia surface of the appliance, the alveolar segments are molded to a more normal position that ultimately facilitates surgical repair of the lip. Once the alveolar segments are approximated, a nasal stent is added to the appliance. The purpose of the nasal stent is to support and shape the dome of the nares and adjacent cartilages. Advantages of presurgical nasoalveolar molding The primary advantages of presurgical nasoalveolar molding addressed in the literature are presented herein [52]. Controlled and predictable positioning of the alveolar segments is achieved without the need

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for lip adhesion surgery or more invasive surgical placement of pin-retained appliances. Most active presurgical orthopedic techniques work under the concepts presented by Matuso et al. [87] in 1984 regarding the high plasticity of cartilage of infants in the first months after birth because of the high levels of hyaluronic acid. Hyaluronic acid is a component of proteoglycan aggregate of the intercellular matrix in the cartilage. Hardingham et al. [55] noted that the increased levels of estrogen that are typically present in neonates increased the level of hyaluronic acid. Haug et al. [60] found that this plasticity is lost in the first few months after birth. It is a reasonable approach to take advantage of the neonates first 3 to 4 months of life for the process of nasoalveolar molding. One surgical procedure may be done for repair of the lip, alveolus, and nose. Presurgical nasoalveolar molding allows the surgeon to perform a gingivoperiosteoplasty and reduces need for extensive tissue dissection for widely separated alveolar segments. Santiago et al. [106] in 1998 presented that in more than 60% of the cases studied, this therapy has reduced the need for secondary alveolar bone grafts in children of mixed dentition. Wood et al. [137] in 1997 presented that midfacial growth is not adversely affected by this therapy. Presurgical nasoalveolar molding has reduced the need for bone grafting of the alveolus in later childhood and reduced the need for early nasal revision surgery.

Conventional obturator prostheses Approximately one fifth of the published articles reported scientific data concerning the management of a conventional obturator prosthesis in maxillectomy patients. The other four fifths represented case reports, review articles, or editorials. In the following discussion these clinical studies and laboratory studies are summarized. Clinical studies about conventional obturator prostheses Surgical obturator prosthesis The importance of a surgical obturator prosthesis, which is placed immediately during the maxillectomy procedure inside the operation room, was proved by a retrospective study [73]. In 23 maxillectomy patients the immediate placement of a surgical obturator prosthesis was compared with the delayed placement in terms of patient outcomes. For 17 patients, a

surgical obturator prosthesis was placed immediately as part of the operative procedure, and for 6 patients, an occlusive dressing only was placed. The group with surgical obturators progressed more quickly and had a more rapid return to normal function. The authors concluded that the immediate surgical obturator prostheses for maxillectomy patients are beneficial. Consequently, the additional efforts and costs are justified for fabrication of an immediate surgical obturator prosthesis. Interim obturator prosthesis The cost effectiveness of interim obturator prostheses was evaluated, which is a sensitive point during this sometimes frustrating treatment period after maxillectomy [65]. A retrospective study analyzed 100 records of patients who underwent maxillectomies. The study quantified the number of patient appointments needed during the interim obturator service period, defined by global 90 days. Of the 100 patients analyzed, 42 had a definitive obturator prosthesis fabricated within 3 months after surgery. On average, 12 appointments (range of 6 – 24 appointments) were recorded for each patient during this 3-month global interim obturator period. The authors concluded that the mean 12 appointments reflect a considerable amount of clinical and laboratory effort and reinforce the national concern of proper reimbursement. Definitive obturator prosthesis A clinical study showed that a well-functioning definitive obturator prosthesis contributed significantly to the improved quality of life of maxillectomy patients by restoring speech and eating function [67]. Forty-seven patients had a maxillectomy with an obturator prosthesis an average of 5.2 years ago, 94% of whom had some of their soft palate resected. The patients were interviewed on the telephone by a trained research interviewer using a series of questionnaires to assess their satisfaction with the functioning of their obturator and psychologic, vocational, family, social, and sexual adjustment. Statistical analysis showed that satisfactory functioning of the obturator prosthesis was found to be the most highly significant predictor of patient’s psychosocial adjustment to the illness and was significantly related to their perception of the negative socioeconomic impact of cancer upon their lives. The most significant predictor of better obturator functioning was the extent of resection of their soft palate and the hard palate. Specific aspects of obturator functioning that most significantly correlated with better adjustments were (1)

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less difficulty in pronouncing words, (2) chewing and swallowing food, and (3) less change in voice quality after surgery. Speech intelligibility after maxillectomy was analyzed in 54 patients with and without obturator prosthesis [123]. With an obturator prosthesis the mean speech intelligibility score of 84% was twice as high than without an obturator prosthesis, which resulted in a mean speech intelligibility score of only 35%. The resection of the anterior portion of the soft palate was one of the factors that influenced the speech intelligibility score of a prosthesis. Speech intelligibility of eight maxillectomy patients was compared with and without obturator prostheses [38]. Oronasal separation and velopharyngeal function were evaluated by use of a specially designed spirometer and endoscope. After placement of maxillary obturator prostheses, only four patients achieved dramatic improvement in speech intelligibility, whereas four patients did not. In the latter, insufficient improvement in speech intelligibility was attributed to velopharyngeal incompetence or unstable prosthesis. Two of three patients with velopharyngeal incompetence did achieve adequate improvement in speech after placement of a speech appliance in combination with maxillary obturator prostheses. An article presented a speech assessment protocol for patients using either obturator prostheses or speech aid prostheses [86]. The protocol was structured according to the executive summary of ‘‘Disability in America: Toward a National Agenda For Prevention,’’ a report formulated by the Institute of Medicine. The results of two patients illustrated that the speech outcome assessment measured the speech intelligibility more accurately for surgically acquired defects caused by cancer of the maxilla or soft palate. Clinical experience suggests that maxillectomy patients who suffer from trismus can present some difficulties in fabricating a conventional definitive obturator prosthesis. Surprisingly, this assumption was not supported in a study that assessed the difficulty in fabrication of a definitive obturator prosthesis in trismus patients [124]. Fifty-four patients who suffered from trismus with less than 20-mm mouth opening were subdivided in four trismus groups: (1) less than 5-mm opening, (2) 5 to 10 mm, (3) 11 to 15 mm, and (4) 16 to 20 mm, which were compared with a control group. The authors reported no apparent differences among the four trismus groups or with the control group. It was suggested that factors other than mouth opening were important in fabrication of an obturator prosthesis, such as a complex anatomy of the resection site, flexibility of

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the lip and cheek tissues, and the scar bundle adjacent to the defect site. The clinical successes of light-cured obturators was evaluated in ten maxillectomy patients [15]. All ten patients had previously worn conventional heatcured hollow bulb acrylic obturators and were currently rehabilitated with light-cured obturators. The obturators were assessed on their performance for weight, retention, speech, eating, and comfort. The authors concluded that patients found light-cured obturators to have a more acceptable clinical performance than conventional acrylic obturators primarily because of a reduction in appliance weight. The study design did not address the sequence of delivered obturator type, however, because no heat-cured acrylic obturator followed a light-cured obturator, which omitted a possible learning curve for patient and dentist. Laboratory studies about conventional obturator prostheses The heat-processed acrylic resin showed its reliability as a ‘‘working horse’’ for maxillofacial prosthetics in a laboratory study. The safest watertight seal between hollow obturator and lid in all test samples was achieved only with the two-flask technique sealed using heat-processed acrylic resin [96]. Encouraging results about visible light-cured resin used for an obturator prosthesis material were published. An in vitro model was designed to determine sorption and permeability of hollow visible lightcured resin obturators in comparison to heat-cured acrylic obturators [14]. The authors found that the visible light-cured resin had a lower sorption value than heat-cured acrylic resin, which made it a more suitable material for producing hollow obturators. Another laboratory study assessed the dimensional change in maxillary prosthetic obturators constructed with three different denture base resins [29]. The results indicated that a visible light-cured resin material had the least percent of change between base and teeth positions. To the knowledge of the authors, however, since the early 1990s no clinical study has been published to support the routine use of visible light-cured resin material as substitute for heat-cured resin material. A conventional definitive obturator is subjected to constant force of gravity beside the intermittent forces encountered during function. The gravity-induced stresses transmitted to the remaining oral structures by various obturator prosthesis framework designs were studied with a photoelastic in vitro model [108]. Frameworks that used I-bar and circumferential

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retainers with buccal retention were most severe, whereas the swinglock and light wire retainers were intermediate in generated stress.

Palatal augmentation prosthesis, palatal lift prosthesis, and speech aid prosthesis Most of the published literature consisted of case reports, review articles, and editorials discussing the treatment of speech problems with maxillofacial prostheses. The few clinical studies that investigated the clinical outcome of these speech appliances are summarized in the following. One study questioned the ability of palatal lift prostheses to stimulate the neuromuscular activity of the velopharynx in evaluating 25 patients who underwent placement of a palatal lift prosthesis for velopharyngeal dysfunction [133]. Nonendoscopic evaluations were audio-videotaped before and after the prosthetic treatment, and the tapes were rated according to three speech pathologists experienced in assessment of patients with velopharyngeal dysfunction. Results of this study neither supported the concept that palatal lift prostheses alter the neuromuscular patterning of the velopharynx nor provided objective documentation of the feasibility of prosthetic reduction for weaning. The challenging results of the latter study were modified by the two following clinical investigations. The effect of a palatal lift prosthesis or of a palatal lift prosthesis with pharyngeal bulb on the levator veli palatini muscle activity during blowing was evaluated by electromyography on eight patients [112]. Electromyography of the levator veli palatini muscle was recorded with a speech appliance in place and then with the speech appliance removed as the subject blew through a tube at three different effort levels. The electromyographic levator activity changed in relation to oral air pressure with either speech appliance in place for all subjects regardless of their speech appliance types. The authors concluded that the effect of a speech appliance to correct velopharyngeal incompetence might consist not only of mechanical obturation of the velopharynx but also of alteration of velopharyngeal function to become similar to normal speakers. The study also summarized the likelihood that the velopharyngeal system could be well regulated so as to exhibit a consistent outcome of velopharyngeal function. Neurologists do not frequently request the use of palatal lift prostheses and palatal augmentation prostheses for dysarthria in patients who suffer from amylotrophic lateral sclerosis, a progressive, adult-

onset neurodegenerative disorder. A retrospective study based on 25 patients who suffered from this disease assessed the efficacy of a palatal lift and augmentation prosthesis on improving speech function and intelligibility through chart reviews, phone interviews, and office interviews [35]. The reported results were encouraging: 84% of the patients treated with a palatal lift prosthesis demonstrated improvement in their dysarthria, specifically in reduction of hypernasality; 76% of the patients benefited at least moderately for 6 months. Of the 10 patients treated with a combination palatal lift and augmentation prosthesis, 6 demonstrated improvement in articulation. Most patients indicated that it was easier to speak with less effort involved when wearing the prosthesis. Consequently, the authors recommended the use of a palatal lift and augmentation prosthesis in amylotrophic lateral sclerosis patients with dysarthria.

Implant-retained maxillofacial prostheses The concept of osseointegration has revolutionized the challenging field of maxillofacial prosthetics and improved patients’ quality of life. Despite this clinical importance, only every fourth publication presented scientific data in form of a laboratory or a human study about intraoral and extraoral implants in maxillofacial prosthetics, which are summarized in the following discussion. Clinical studies about intraorally placed implants in maxillofacial prosthetics The success rate of intraorally placed implants in irradiated bone is described by three studies. A preliminary report described the effect of hyperbaric oxygen therapy on the osseointegration of titanium implants in irradiated bone [122]. Preoperatively and postoperatively, hyperbaric oxygen therapy was administered to four patients at the level of 2 or 3 atmosphere absolute. Of the 21 fixtures placed, only 1 was lost because of lack of osseointegration, which resulted in a surprisingly high survival rate of 92%. Another research group published less favorable results about implants that retain a maxillary obturator in two clinical retrospective studies [81,104]. In the first retrospective study, 23 patients received a total of 85 osseous integrated implants used for retaining maxillary obturators between 1985 and 1993 [81]. Implants were placed at the time of ablation or subsequently were uncovered after a

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6-month period of integration and were used in prosthetic rehabilitation. Thirteen of the 23 patients with 50 implants were radiated with a dose ranging from 5040 to 7940 cGy. The authors identified radiation therapy and the specific maxillary implant region as risk factors: (1) radiation therapy reduced the success rate for implants from 80% to 55%, and (2) in the anterior region of the maxilla an implant success rate of 86% was determined in comparison to only 57% in the posterior region of the maxilla. All implant failures occurred at the second stage in this clinical study. A second study with more recent data concerning the same clinical question reported more favorable results for irradiated maxillectomy patients [104]. Twenty-six patients were included with 102 implants placed, from which there were 19 intact withdrawals because of implant loss caused by recurrent disease or patient death, 5 implants with unknown status, 24 implant failures, and 54 functional implants. In this patient population, the overall success rate for implants was 63% in the irradiated group. Interestingly, the success rate for implants inserted before radiation was 67%, better than for implants placed after radiation, with only 50%. The success rate for implants in the nonirradiated group was 82.6% which confirmed the results of Lorant et al. [81]. The study linked low success rates with (1) implants placed during tumor resection, (2) implants placed within the maxillectomy defects, and (3) implants that received postoperative radiation. The authors stated that the success rate of 50% to 67% for implants placed intraorally in the irradiated maxillectomy patient is not overwhelming but is still no reason not to apply them. Most implant failures, with 18 of 24 (75%), occurred either at stage II surgery or before loading. This observation raises the question whether the interim obturator prosthesis might be responsible for extended transmucosal loads or if an extended healing time in the irradiated group might have improved the implant success rate. Clinical studies about extraorally placed implants in maxillofacial prostheses The placement of extraoral implants is important for improving a patient’s quality of life because conventional prostheses may lack adequate retention and stability, which diminishes the patient’s confidence that the prosthesis will remain in place during routine activities. One study assessed the role of extraoral implants in improvement in quality of life by a questionnaire [9]. Five patients with acquired midface defects were treated with 19 titanium endosseous root-form implants to provide retention and stability for prostheses. Patients responded to a ques-

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tionnaire rating overall use, effectiveness, and satisfaction of their prosthesis before and after the use of implants. Analysis of the questionnaire indicated an improvement in the quality of life for the patients with an implant-retained prosthesis. The same study also reported 3 (18%) failed implants out of 17 prosthetically used implants. The possibility of achieving osseointegration around an orbital defect was not as good as in the mastoid process. Higher success rates were reported for implants placed in the mastoid region in comparison to the orbital region, especially after radiation [136]. The same observation was made by evaluating osseointegrated implants in the restorative treatment of auricular and orbital defects over a 5-year period [62]. The total success rate for survived implants was 95.6% in the auricular defects and only 67.2% in the orbital defects. More promising results of successful implantretained orbital prostheses in patients treated after orbital exenteration were reported. Successful osseointegration was achieved in 17 (94.4%) of 18 fixtures in five nonirradiated patients; however, one patient who underwent irradiation lost all 4 fixtures that had been placed [90]. Only five patients were evaluated after an average follow-up of 16 months. Similar results of implant-retained orbital prostheses were reported for nine patients with 25 implants [70]. In seven cases with 19 implants, orbital prostheses were fabricated and have been worn for an average of 3 years. Only 1 of the 19 implants was lost after 3 years because of load stress. A surprisingly good outcome was reported for the same anatomic region in irradiated cancer patients. A retrospective study evaluated 24 implants that were placed in six patients for prosthesis fabrication with a history of orbital exenteration and irradiation for oncologic tumors of the head and neck [68]. Twenty-one implants were ultimately used to retain six orbital prostheses. Two implants failed to maintain osseointegration during the follow-up period and subsequently were removed without complications, which led to an overall integration success rate of 90.5% over a mean follow-up period of 32.8 months. These reported optimistic results for the orbital region are in disagreement with a clinical study over a 7-year period of 23 craniofacial implants placed in eight irradiated and nonirradiated orbital detects [92]. Implants placed in the orbital region demonstrated a high failure rate: implant success rate was only 35% with 7 intact implants out of 20. Implant success rate in the nonirradiated patients was 37.5% (3/8), and the success rate for radiated patients was only 33.3% (4/12). Most implant failures occurred late as opposed to early in the study period.

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A few studies addressed the importance of periimplant hygiene in maxillofacial patients. The publication of Jacobsson et al. [62] found some skin problems in approximately 10% of the patients treated with auricular and orbital implants, whereas the remaining 90% of the patients showed no or minimal problems. One publication studied the crevicular microflora activity by collecting samples for crevicular sites surrounding 17 craniofacial implants [116]. The results demonstrated the presence of opportunistic pathogens in many sites, surprisingly regardless of the subject’s hygiene efforts. In another study, the importance of periimplant hygiene was stressed [46]. Two of seven patients with percutaneous craniofacial implants exhibited adverse skin reactions that were associated with heavy sebaceous crusting, skin cultures positive for Staphylococcus aureus, higher Periotest values, and thicker periabutment tissue with higher mobility. It was determined that these factors can predispose the patient to local infection, which, if ignored, can result in failure of the implant. The authors concluded that adequate patient hygiene is crucial to maintaining healthy tissues in the periimplant abutment site. The implant success rate and the soft tissue responses of osseointegrated implants that support orbital prostheses were recorded [92]. Evaluations were conducted at 6-month intervals, and for the study period there were 80 visits/sites. Implants placed in the orbital region demonstrated common soft tissue responses and a high failure rate. The study revealed that only 42% of the visits/sites demonstrated an absence of inflammation; however, a high percentage (58%) showed some sign of tissue response around the orbital implants: 24% of visits/sites demonstrated slight redness, 14% demonstrated periimplant red and moist tissues, 6% demonstrated granulation tissue associated with the implants, and 14% demonstrated infection of the periimplant soft tissues. Implant success rate was only 35%. Implant success rate in the nonirradiated patients was only 37.5%, and the success rate for radiated patients was still 33.3%. Most implant failures occurred late as opposed to early in the study period. The authors concluded that orbital implants require meticulous hygiene maintenance and should be placed in patients who understand that long-term success rates may be low. Laboratory studies about implants in maxillofacial prostheses The retention force between either cylindrical or telescopic magnetic inserts with secondary magnets was measured at different angles [91]. The mean

force in the axial direction was 1.8 N. An increasing angle between force direction and insert-axis decreased the retention force, which stresses the importance of the optimal positioning of implants to retain facial prostheses. Another in vitro study evaluated the mechanical behavior of retention systems with a custom-designed loading/measuring apparatus that represented a typical auricular situation with three points of retention [31]. The test apparatus provided vertical and horizontal loads in five locations. The system was used to test two ball-and-socket attachments (Dalla Bona, NobelBiocare (Yorba Linda, CA)), cast and performed bar and clips (NobelBiocare (Yorba Linda, CA)), and three magnet systems (Dynamag, Neomag, Technovent). The loading/measuring apparatus also was used to evaluate the performance of two facial prosthetic adhesives. Not surprisingly, the retention systems used in craniofacial osseointegration offer more predictable retention than the facial prosthetic adhesives. The researchers concluded that the mechanical systems are best suited to situations in which tensile and shear forces exist. Magnet systems were suggested to be applied only where tensile forces are anticipated or where horizontal forces on the implants are to be avoided.

Economic aspects As already observed for other fields in maxillofacial prosthetics, only a few articles studied scientifically the issue of economic aspects, whereas most publications included case reports, editorials, and review articles. Maxillofacial prostheses can contribute much to a patient’s quality of life; however, the availability of an adequate service varies internationally. In most major cities in the United States, there is at least one adequate-to-superior medical facility that offers such patients a substantial choice of treatments. This may not be the case in other countries, however. One United Kingdom survey of oral and maxillofacial surgeons reported that maxillectomies were performed by 83% of the surgeons surveyed (most surgeons perform one to five cases per year) and that 38% of the surgeons do reconstructive surgery, but only in 10% of the cases [3]. Only 65% of the surgeons have access to the services of a restorative dentist, and this did influence 19% of the surgeons’ decisions about whether to reconstruct surgically or prosthetically [3]. Access to and knowledge of maxillofacial prosthodontists and their services are at issue everywhere, and communication between surgeons and prosthodontists is neces-

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sary for all treatment options and choices to be provided to patients. In Scotland, the provision and requirements for adequate hospital maxillofacial prosthetic services were examined by a questionnaire of 90 oral plastic and ear, nose, and throat surgeons [10]. Information was collected on numbers and types of cases seen, prosthetic treatment provided, and involvement of consultants in restorative dentistry in their management. The results showed that most surgeons only had treated a small number of such patients. The authors stated that that result led a dilution of experience, and they were in favor of supraregional specialist centers, which facilitate the needed improvement of multidisciplinary planning clinics for such cases. In the United States, a questionnaire study was conducted in 1991 regarding the frequency of maxillofacial procedures and insurance coverage [131]. A total of 342 survey instruments were received from a total of 690 mailed, which represented a total of 18,410 maxillofacial procedures performed by the survey population. Members of the American Academy of Maxillofacial Prosthetics performed procedures at a significantly higher rate than did members of the American College of Prosthodontists. A greater procedure rate was observed for (1) prosthodontists in the South, Midwest, and Southwestern regions, (2) the hospital setting, with 2 plus 1 year additional postgraduate maxillofacial training, and (3) the 45-to 54-year-old age group. Insurance covered most maxillofacial procedures but was not uniformly distributed within predictor variables nor between procedures. A study by the same research group published a preliminary analysis that evaluated the Medicare Resource Based Relative Value Scale as it relates to maxillofacial prosthetics [130]. The degree of difficulty and treatment time were determined for nine maxillofacial procedures codified in Physicians’ Current Procedural Terminology, edition 4. A preliminary survey indicated that practice expense and malpractice expense ratios were 66% and 1%, respectively. Unfortunately, a correlation of maxillofacial procedures to other medical procedures could not be made because an additional study with a greater sample size is needed in the near future. The results of a national survey of members of the American Academy of Maxillofacial Prosthetics were reported 2 years later based on the estimated relative value of provider work for nine existing and two proposed coded maxillofacial prosthetic services [27]. The work estimated a necessary component of a larger effort by the Health Care Financing Administration to establish a resource-based relative value scale for reimbursement for all Medicare services. Analysis of the survey

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data revealed agreement regarding the relative value of work for maxilllofacial services compared with other established medical procedures and that the median values for each service were acceptable initial estimates.

Near future of maxillofacial prosthetics Undoubtedly the routine use of osseointegrated implants in the remaining midfacial skeleton has greatly facilitated prosthetic rehabilitation of large maxillary defects [81]. The relatively low success rate of implants in the orbital region and in irradiated bone regions remains a challenge in the future, however. From an economic standpoint there is lack of routine financial coverage by third parties, especially for state-of-the-art treatments with implants. At least in Europe, the improper insurance coverage of maxillofacial prosthetics in cancer patients often harms a patient’s quality of life, leaving one to hope for a health system without bureaucratic overregulations in the near future.

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Wound closure materials Robert Gassner MD, DMD, PhD* Department of Oral and Maxillofacial Surgery, University of Innsbruck, Innsbruck, Austria Department of Oral and Maxillofacial Surgery, University of Pittsburgh, Pittsburgh, PA, USA

Proper wound closure is part of successful overall wound care after the assessment of patient and sustained wound, anesthesia, debridement, and irrigation of the wound. Postoperative care enhances excellent outcome or reduces impaired healing situations. Wound healing disturbances mostly result from neglecting one or more of these steps in wound care regardless of the kind of wound closure material used. The principles of wound care have remained remarkably the same over the years. Although most lacerations heal without sequelae regardless of management, mismanagement may result in wound infections, prolonged convalescence, unsightly and dysfunctional scars, and, rarely, mortality [29,66]. Wounds are one of the most commonly encountered problems in the emergency department. They rival respiratory tract infections as the most common reason people seek medical care [44,55]. Five years ago almost 11 million wounds were treated in emergency departments throughout the United States. At an average charge of $200 per patient, this translates to more than $2 billion annually. When nonemergency or elective incisions are included, approximately 90 million skin-suturing procedures are performed each year [70]. The goals of wound management are simple: avoid infection and achieve a functional and esthetically pleasing scar [66]. These goals are achieved by reducing tissue contamination, debriding devitalized tissue, restoring perfusion in poorly perfused wounds, and establishing a wellapproximated closure.

* Department of Oral and Maxillofacial Surgery, University of Pittsburgh, G-33 Salk Hall, Pittsburgh, PA 15261, USA. E-mail address: [email protected] (R. Gassner).

Most lacerations require primary closure. Primary closure results in more rapid healing and reduced patient discomfort than does secondary closure. The most commonly used method for closing lacerations is suturing. Most wound care practices are empirical or based on animal wound models; few are based on well-designed clinical trials. Most studies of laceration management have focused on the wound infection rate as the primary outcome, despite the fact that wound infections are relatively uncommon (less than 5% of lacerations) [30]. Although all traumatic lacerations should be considered contaminated, most have low bacterial counts (fewer than 100 organisms per gram of tissue), well below the infectious inoculum of 105 or more organisms per gram [59]. Most infected lacerations heal without complications other than the occasional unsightly scar. Patients are most concerned with the cosmetic appearance of their healed lacerations [54], and the focus of physician’s and dentist’s wound research is shifted toward measuring wound cosmesis as the primary outcome, neglecting pressure on development of new biomaterials for wound closure.

History of wounds Humans have managed wounds from the beginning of civilization. The earliest reports of the use of artificial materials come from the Edgar Smith papyrus, which details the use of sutures and wound closure devices around 4000 BC [50]. Initial treatments for wounds consisted of herbal balms or draughts, with application of leaves or grasses as bandages [20]. Ointments were made from a wide variety of animal, vegetable, and mineral substances. Wounds were mostly left open, although wound

1042-3699/02/$ – see front matter D 2002, Elsevier Science (USA). All rights reserved. PII: S 1 0 4 2 - 3 6 9 9 ( 0 2 ) 0 0 0 0 9 - 2

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closure using the jaws of ants was used by some cultures [74]. The world’s oldest suture was placed by an embalmer on the abdomen of a mummy in approximately 1100 BC [39]. During the Middle Ages, pus was believed to be necessary for healing; as a result, various agents were used to promote suppuration. Advances in the fields of anesthesiology and surgery during the past two centuries have led to the development of many of the practices that are currently prevalent [29,66]. These advances are based on thorough debridement and cleansing of wounds and the use of aseptic wound closure techniques. Only recently has wound care been investigated systematically in the laboratory and in clinical arenas.

Epidemiology of wounds Approximately one third of lacerations occur in adults between the ages of 19 and 35 years [30], predominantly of male gender. Most wounds are found either on the head or neck (50%). The most frequent mechanism of injury is application of a blunt force, such as a falling against obstacles. Other agents of injury include sharp instruments, glass, and wooden objects [30]. Although mammalian and human bites continue to receive much attention, they are a relatively rare cause of wounds.

A detailed history of allergies to any agents is essential (e.g., anesthetic agents and antibiotics, latex products). Tetanus immunization status should be verified.

Wound assessment Besides discussion on the necessity of sterile conditions for treatment of wounds [6], powder-free gloves reduce the risk of any foreign body reactions or infections that theoretically may result from the introduction of talc particles into the wound. The wound is examined meticulously in all cases. Proper lighting and control of bleeding are required to identify any injury to vital structures (such as nerves and tendons) and foreign bodies that may contaminate and lead to chronic infection and delayed healing [29,66]. Failure to diagnose foreign bodies is the fifth leading cause of litigation against emergency physicians [29]. Other common wound-related causes of litigation include the development of wound infections and missed injuries of tendons and nerves. Crush injuries, which tend to cause greater devitalization of tissue, are more susceptible to infection than are wounds that result from the more common shearing forces [10].

Anesthesia of the wound Patient assessment Increased wound infection rates or delayed healing exists for patients with diabetes mellitus, obesity, malnutrition, chronic renal failure, advanced age, and use of steroids [12]. All of these risk factors, together with the use of chemotherapeutic agents and other immunosuppressive agents, may delay wound healing by affecting inflammation and the synthesis of new wound matrix and collagen [31]. Healing also may be impaired in inherited and acquired connective-tissue disorders, such as EhlersDanlos syndrome, Marfan syndrome, osteogenesis imperfecta, and protein and vitamin C deficiencies. The tendency of the patient to form keloids should be ascertained because this may result in a poor scar [66]. Keloids extend beyond the boundaries of the original injury and are largely determined by genetic or racial predisposition. Conversely, hypertrophic scars, which remain within the boundaries of the original injury, usually result from a tissue deficiency or from the fact that the wounds are not parallel to the lines of minimal skin tension [66].

For adequate evaluation and management before wound closure, many lacerations require anesthesia. There are two major classes of local anesthesia: esters and amides. Several possibilities for patientfriendly administration of local anesthesia exist, including warming of the anesthetic solution to body temperature, slowing the rate of injection, and injecting the local anesthesia through the wound edges of the laceration instead of through the intact surrounding skin. Use of smaller needles and subcutaneous rather than intradermal injection also has been suggested to result in less pain [35,63]. Buffering of the local anesthesia with sodium bicarbonate at a ratio of 1:10 provides a more rapid and less painful onset of anesthesia.

Wound preparation, debridement, and irrigation Cleaning the wound is an important step before wound closure. Although scrubbing with a highporosity sponge in highly contaminated wounds achieves beneficial effects, such as bacteria and

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particulate matter removal, deleterious effects may occur caused by intense scrubbing that results in further tissue damage. Nonviable tissue may impair the ability of the laceration to resist infection. Surgical debridement of any crushed or devitalized tissue is one of the most fundamental aspects of wound preparation [25,66]. Eyebrow hair should not be removed because it may result in abnormal regrowth. Removal of other hair surrounding a laceration may help facilitate meticulous wound closure. Because many bacteria normally reside in hair follicles, shaving of the hair before repair may increase wound infection rates [66]. Devitalized tissue such as fat, muscle, and skin that remains further impairs the ability to resist infection [25]. Removing such tissue mechanically and surgically is an essential part of wound management. Mechanical debridement may be performed by surgical excision, scrubbing with a surgical sponge, or high-pressure irrigation [66]. Considerable debate exists regarding the exact methods of irrigation, especially on irrigation impact pressures and the nature of the irrigant solutions [25,29,66]. Normal saline solution remains the most cost-effective and readily available choice [15,28]. Because of their tissue toxicity, detergents, hydrogen peroxide, and concentrated forms of povidone-iodine should not be used to irrigate wounds [15,28]. In highly vascularized areas that contain loose areolar tissue, such as the eyelid, high pressures should be avoided [45]. Irrigation may not be required for all low-risk wounds in the face [28].

Wound closure More than 40 different types of sutures exist to close wounds and incisions [66]. It seems to be irrelevant how the wound is held together as long as good healing and esthetics result [41,61]. The time during which wound closure is safe must be tailored individually on the basis of causation, location, and host factors. When wounds are not closed because of a high risk of infection, delayed primary closure should be considered after 3 to 5 days, when the risk of infection decreases, especially if they are large, may have a poor cosmetic outcome, or are associated with discomfort or inconvenience [66]. Most wounds should be closed primarily to reduce patient discomfort and speed healing. Wounds at low risk for infection can be closed 12 to 24 hours after the injury, but for wounds at high risk (contaminated wounds, those in locations with poor vascular supply, and those in immunocompromised patients), primary

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closure should take place within approximately 6 hours. There is a direct relation between the time from the injury to closure of the laceration and the risk of subsequent infection, but the length of this ‘‘golden period’’ is highly variable [5]. Each individual wound must be considered separately, taking the time from the injury until presentation into account in addition to laceration location, contamination, risk of infection, and importance of cosmetic appearance before deciding the form of wound closure. Wounds that are not closed primarily because of a high risk of infection should be considered for delayed primary closure after 3 to 5 days, when the risk of infection decreases.

Options for wound closure The ideal wound closure material permits a precise wound closure, is easily and rapidly applied, is painless, is inexpensive and of low risk to patients and providing persons, and results in minimal scarring with a low infection rate. Sutures Sutures are the most commonly used wound closure material. Nonabsorbable sutures, such as nylon and polypropylene, retain most of their tensile strength for longer than 60 days, are relatively nonreactive, and are appropriate for closure of the outermost layer of the laceration [42,57,71]. Removal of nonabsorbable sutures is required. Absorbable sutures are generally used for closure of structures deeper than the epidermis. In general, synthetic absorbable sutures are less reactive and have greater tensile strength than sutures from natural sources, such as catgut. They increase the time during which the healing wound retains 50% of its tensile strength from less than 1 week to as long as 2 months. Chromic gut lasts for up to 2 weeks and is associated with tissue reactivity. Polyglactin and polyglycolic acid maintain tensile strength for 20 to 28 days and are associated with minimal tissue reactivity. Some synthetic absorbable sutures (e.g., polydioxanone, polyglyconate) preserve their tensile strength for as long as 2 months, which makes them useful in areas with high dynamic and static tension. Use of these sutures should be limited to deeper structures because they may become extruded over time (Table 1) [42,57,71]. Although use of absorbable sutures is generally reserved for subcuticular tissues, rapidly dissolving forms may be used to close the wounds in children

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Table 1 Characteristics of absorbable sutures Suture material

Wound tensile strength

Knot security

Tissue reactivity

Security (d)

Polyglactin (Vicryl) Polyglycolic acid (Dexon) Polyglyconate (Maxon) Polydioxanone (PDS) Chromic gut Surgical gut

Good Good Excellent Excellent Average Average

Good Excellent Average Average Average Poor

Low Low Least Least High High

30 30 45 – 60 45 – 60 10 – 14 5–7

and thereby avoid the discomfort associated with suture removal [29,66,68]. Generally, synthetic and monofilament sutures are preferred over natural and braided sutures because they result in lower rates of infection (Table 2) [29,66]. Staples Staples can be applied more rapidly than sutures. They are associated with a lower rate of foreign body reaction and infection [58]. In general, staples are considered particularly useful for scalp, trunk, and extremity wounds and when saving time is essential (e.g., mass casualties, patients with multiple trauma wounds) [7,58]. They do not allow as precise a closure as sutures, however, and are slightly more painful to

remove. In animal models, staples are associated with lower rates of bacterial growth and lower infection rates than sutures [7]. In clinical series, these effects may be statistically significant but are of limited clinical significance [27]. Comparing suture and staple repair of simple pediatric scalp lacerations stapling resulted in shorter wound closure times, shorter overall times for wound care and closure, and fewer expenses in terms of equipment and total cost based on equipment and physician time [34]. Adhesive tapes Advocates of adhesive tapes proclaim the superiority of tape wound closure because of the resistance to infection of the underlying wound that is greater than

Table 2 Pros and cons of common wound closure materials Technique

Advantages

Disadvantages

Suture

Classic method Careful closure Most resilient tensile strength Lowest dehiscence rate Risk of needle stick

Requires removal Requires anesthesia Greatest tissue reactivity Most expensive method Slowest method

Staples

Rapid application Low tissue reactivity Low cost Low risk of needle stick

Sloppy closure Interferes with older generation imaging technology (CT, MR image)

Tissue adhesives

Rapid application Patient comfort Resistant to infection No need for removal Low cost No risk for needle stick

Less tensile strength than sutures Dehiscence in high-tension areas (such as joints) Cannot be used on hands No bathing or swimming

Surgical tapes

Rapid application Patient comfort Lowest tissue reactivity Lowest infection rates Low cost No risk for needle stick

Frequently falls off Less tensile strength than sutures Highest rate of dehiscence Includes toxic components Cannot be used in the hair Cannot be wetted

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in wounds that contain sutures, staples, or tissue adhesives [16,51]. This unique resistance to infection of tape-closed wounds is attributed to the absence of the sutural, tissue adhesive, or staple foreign bodies in the wound. These wounds require the use of adhesive adjuncts (e.g., tincture of benzoin) that increase local induration and wound infection. Because adhesive adjuncts are toxic to wounds, care should be taken that they do not access the wound. Although the various surgical tapes have different degrees of adhesion, porosity, breaking strength, and elasticity, tapes alone cannot maintain wound integrity in areas subject to tension [60,62]. They have been used in an estimated 1 billion patients but are more often used after suture removal to decrease tension on the wound until they fall off [66]. Tape wound closure is preferred in children because of low discomfort during application. Tissue adhesives The skin of wounds also can be closed by tissue adhesives. They should be applied only topically. Tissue adhesives cannot be used intraorally and cannot replace deep sutures. Approximately 30% of laceration repairs and closure of many surgical incisions may be amenable to skin closure with tissue adhesives; thus, they are not a replacement for sutures but do offer an alternative [55]. Modern tissue adhesives contain cyanoacrylates. Cyanoacrylates were first synthesized in 1949 and used clinically in 1959 as agents to glue skin wounds [29]. Monomeric cyanoacrylates polymerize in the presence of hydroxyl ions, which can be found in water and blood, thereby bonding with the skin. The ethylene portion of the molecule polymerizes. The initial derivatives were methyl-2- and ethyl-2-cyanoacrylates. These derivatives were effective, but the shorter alkyl chains degraded rapidly into cyanoacetate and formaldehyde. These products had significant tissue toxicity that resulted in acute and chronic inflammation. With longer alkyl chains, the toxicity decreases as a result of slower degradation, which limits the accumulation of byproducts. N-butylcyanoacrylates are less toxic than shorter-chain cyanoacrylates and maintain a stronger bond. Longer-chain N-2-butylcyanoacrylate adhesive has been used for wound closure in Canada and Europe for more than 20 years. Throughout this time no adverse effects have been reported [36]. A study in 1985 reported that isobutyl-2-cyanoacrylate implanted into a rat peritoneum induced sarcomas. This study was conducted in rats prone to sarcomas and was never duplicated; nevertheless, it did result in a lack of Food and Drug Administration approval for

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use in the United States [73]. Butylcyanoacrylates do have limitations. After polymerization, the adhesive is brittle and can fragment if flexed. Because of these limitations, butyl compounds have been used for areas that do not cross wound creases and for shorter wounds [29,55]. 2-Octylcyanoacrylate, formulated with plasticizers, is even more stable, has greater flexibility, and maintains a stronger bond. It degrades much more slowly, which leads to its classification as nontoxic. 2-Octylcyanoacrylates (e.g., Dermabond, Ethicon, New Brunswick, NJ) were approved for use in the United States in 1998. Within the first month of availability, more than 3 million units of Dermabond were ordered [29,55]. Studies have shown wound breaking strength to be equal to that of suture-repaired wounds at 5 to 7 days, but the day 1 breaking strength is only approximately 10% to 15% that of a wound closed with 5-0 monofilament sutures [47,76]. Wounds are evaluated and cleaned in a standard fashion. If the wound edges are traumatized or involve multiple tissue layers, appropriate debridement or buried sutures or both are necessary. For more superficial wounds, no buried sutures are needed. The wound edges are manually approximated with fingers or forceps while care is taken not to apply the adhesive between the edges. Interposing the glue into the wound results in greater scarring. The wound is held in position for 30 seconds to complete polymerization. Besides displaying three-dimensional tensile strength, the tissue adhesives act as their own dressing. The wound is waterproof and can be wet in a shower, although soaking may result in decreased strength or peeling of the dressing. Experimentally they have antimicrobial effects against gram-positive organisms and the potential to decrease the rate of wound infections [52,53]. The adhesive holds well on the face, and it usually stays on for 7 to 14 days then sloughs off with the epidermis. There is no need for follow-up for suture removal, which makes this method of care attractive. In general, cyanoacrylates are less expensive than sutures and staples, and patients prefer them [48]. Using tissue adhesives for wound repair is a manual skill, like suturing, and requires careful application. Being a topical closure, wound healing is impaired when adhesive gets between the wound edges, which prevents epithelialization and promotes foreign-body reactions [18,72]. When used properly for topical wound closure, tissue adhesives result in fewer foreign-body reactions than sutures do and can decrease infection rates in contaminated wounds [52]. Histologic studies report no differences in wound healing characteristics between sutured and tissue adhesive repaired wounds [32,47,52].

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Prospective, randomized trials comparing Histoacryl Blue (a butyl-2-cyanoacrylate from Braun, Germany) with 5-0 or 6-0 sutures for repair of small facial lacerations revealed equivalent cosmetic outcome at 3 months and at 1 year [29,56,65,66]. The assessment of wounds 3 months after injury and wound repair provided a good measure of long-term cosmetic outcome [56]. The short-term infection rate was slightly higher for Histoacryl Blue but not statistically different, and the wound dehiscence rate was statistically equivalent. The time to wound closure was more than 50% shorter for the group treated with 2-octylcyanoacrylate. Subset analysis of the data suggested that small lacerations aligned against lines of minimal tension may benefit most from the use of tissue adhesive rather than sutures [65]. Octylcyanoacrylate provides the greatest threedimensional tensile strength of all the cyanoacrylates and is a needleless alternative to sutures for closure of most facial lacerations. It provides excellent cosmetic appearance comparable to that achieved with sutures [4,43,49,56]. If lacerations cannot be approximated manually and skin edges cannot be held together without a lot of tension, the use of tissue adhesives is inappropriate. Adequate strength to the wound closure requires three or four coats of 2-octylcyanoacrylate, which is associated with heat release during polymerization as an exothermic reaction. When tissue adhesives result in suboptimal wound closure or must be removed for some other reason, bathing or application of antibiotic ointment or petroleum jelly may accelerate removal. Acetone can be used when more rapid removal is necessary [29,66]. Opponents of tissue adhesives for wound repair note that adhesives seem to act as a barrier between the growing edge of the incision [16,17,41]. This barrier prevents intimate apposition of the wound edges and delays wound healing. The breaking strength of wounds closed with adhesives is believed to be significantly lower than that of taped wounds without the adhesive. Tissue adhesives also potentiate the development of infection in contaminated wounds. Although this technique may be efficient, the use of tissue adhesives—nonabsorbable compounds whose long-term toxicity has not been proved efficacious—is an appropriate substitute for suturing wounds [16,17,41]. If the cyanoacrylate adhesive gains access to the wound, it remains there despite an increased resistance to wound infection. When wounds that contain tissue adhesives become infected, magnification loupes are necessary to identify and remove the adhesive deposits from the wound. Design and development of absorbable adhesives are of utmost importance to allow wound repair

without infection and with the most esthetically pleasing scar. When comparing 2-octylcyanoacrylate and nonabsorbable sutures or staple repair for wounds in children, the tissue adhesive 2-octylcyanoacrylate was revealed to be an acceptable alternative to conventional methods of wound repair with comparable cosmetic outcome [8]. The use of tissue adhesives has been reported in blepharoplasty [24] and has emerged in the management of corneal trauma [38]. Cyanoacrylates also have been used experimentally in skin, bone, and cartilage grafting, corneal and eyelid surgery, and occlusion of esophageal varices and cerebrospinal fluid leaks [29,66]. Autologous fibrin tissue adhesive made from the patient’s own blood and commercial fibrin sealants might offer a possibility for treatment of intraoral lacerations and wounds but require further examination and proof. Fibrin tissue glue proposes to be an alternative mode of managing hemostasis and wound healing. There is much inconsistency in the data secondary to the use of various fibrin sealant preparations, different animal models and clinical situations, and different application techniques. A consequence to this is the likelihood that different fibrin sealant preparations may be preferred for different clinical situations [11,40]. Comparing fibrin glue, as a bioadhesive, with traditional sutures in closing mucosa over exposed mandibular plates in a cat-model revealed fibrin glue to be safe and well tolerated in cats. Glue application required a shorter operative time and was associated with fewer occurrences of granulation and ulceration when compared with suture fixation. Further studies are indicated to titrate the concentration of fibrin [21]. A prospective, randomized, blinded study using fibrin sealant at the wound site revealed significantly reduced pain and decreased chance of experiencing emesis after tonsillectomy [23]. Wound healing might be enhanced by combining platelet concentrates and fibrin adhesive [67]. Dressings Nonadherent wound dressing for at least 24 to 48 hours has a shielding effect until enough epithelialization is present to protect the wound from gross contamination [64]. Maintaning a moist environment around the wound also has been shown to speed the rate of epithelialization [26,75]. For burns in which the traditional split-thickness skin graft is not available, coverage of open wounds is the fundamental problem. Topical antibiotic dressings and porcine skin have been used extensively [46]. The use of

R. Gassner / Oral Maxillofacial Surg Clin N Am 14 (2002) 95–104

xenografts was limited as a temporizing measure, because of intense foreign body reactions and eventual rejections accompanied by a high rate of graft infection. New generations of synthetic skin equivalents have been shown to be superior, with a considerable reduction in the partial-thickness burn healing rate and hospitalization in a clinical trial [14]. These skin equivalents have a bilayered structure. Examples include Apligraf (Novartis, East Hanover, NJ), with live keratinocytes on an acellular dermal matrix, and Integra (Integra LifeScience, Plainsboro, NJ), which consists of reconstituted collagen and danaparoid sodium (chondroitan sulfate) backed by a polymer layer. Both products are currently in use as biologic dressings for the coverage of burn wounds [19,46]. Postoperative care In general, decontamination is far more important than using antibiotics. Antibiotics should be reserved for human and animal bites and for intraoral lacerations and open fractures [13,69]. Avoiding exposure of the wound to sun reduces the likelihood of complications such as hyperpigmentation [66]. Sutures or staples should be removed after approximately 7 days. Periorbitally placed sutures should be removed sooner (within 3 – 5 days) to avoid formation of unsightly sinus tracts. Sutured or stapled wounds should be kept clean and gently cleansed after 24 to 48 hours. Patients with tissue adhesives in place may shower, but they should avoid bathing and swimming. Prolonged moisture loosens the adhesive bond. Gentle blotting to dry the area is preferred to repeated wiping [3,29]. Elevation of the injured area decreases edema formation. Patients should be instructed to observe the wound for erythema, warmth, swelling, and drainage because these findings may indicate infection. Use of standardized wound care instructions improves patients compliance and understanding.

Future prospects Wound healing is no longer considered to consist of three distinctive phases (inflammation, proliferation, and remodeling). It is a dynamic process with overlapping sequences of coordinated cellular events that involve multiple cell types and a cascade of cytokines and growth factors that control the processes of cell migration, proliferation, matrix synthesis, and turnover [33,46]. New biomaterials will emerge as structures that combine several functions within the same device

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and display mechanical support function and sitespecific drug delivery release for application of growth factors and antibiotics to the wound. Finally, biodegradation of the polymer matrix requires the design of custom-made materials with a range of material properties [37]. A way of research is directed toward modifying existing nondegradable polymers into biodegradable, photodegradable, or hydrolyzable polymers by chemical alteration or inclusion of additives (e.g., sensitizers and biopolymers) and developing methods to exploit native biodegradable polymers [2]. Besides directly biodegradable materials [poly(lactic acid), poly(caprolactone), poly(glycolic acid), and related polymers], hydrolyzable polymers (e.g., polyesters, polyanhydrides, and polycarbonates) and modified biopolymers (cellulose, starch) are of particular potential [2]. Currently, although multiple growth factors have been tested clinically, the only growth factor approved by the Food and Drug Administration for clinical use is recombinant human platelet-derived growth factor-BB. This growth factor was used in diabetic foot ulcers and culminated in three pivotal trials that involved more than 1000 patients, who demonstrated consistent 10% improvement over controls in the complete healing of ulcers [46]. Marketed as 0.01% Regranex gel (becaplermin), plateletderived growth factor is currently only approved for the indication of diabetic ulcers, although other studies are ongoing [22]. The variability in healing and the multiple factors that impair healing (ischemia, bacteria, aging, and suboptimal nutrition) undoubtedly explain much of the difficulty in demonstrating a therapeutic effect with growth factors [46]. Laser-assisted skin closure was shown to improve the wound healing process in male hairless rats when compared to control wounds closed with conventional suture techniques. Laser-assisted skin closure led to an acceleration and improvement of wound healing with earlier continuous epidermis and dermis and a thinner resulting scar [9]. Clinical application of biomaterials research will steadily increase in the future of wound closure procedures [50]. Because currently used biomaterials for wound closure display proinflammatory signs and foreign body reaction instead of sound biodegradation, new biodegradable polymers will emerge and change the field of currently widespread and wellaccepted biomaterials such as poly(glycolic acid) and poly(lactic acid). A promising field is nontoxic biodegradable peptide-based polyurethanes with carbohydrate-sidechains as soft extender [1,78]. They are superior to most existing polymers because of

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significant higher biocompatibility that allows cell growth of various types. They do not display noticeable foreign body reactions, antibody responses, inflammatory reactions, or necrosis of the surrounding tissue and permit covalent binding of substrates to be released as bioactive forms to the wound out of sutures, adhesives, and wound dressings. Currently, conventional nonbiodegrading polyurethanes are used extensively in biomedical applications [1,78]. Future advances in wound closure will involve the development of improved suture biomaterials and wound closure tapes, absorbable adhesives, and wound dressings [16]. Nontoxic absorbable materials will revolutionize wound care. Closure of wounds without the need for sutures will be a major advancement, an opportunity to improve care for patients, especially children, and reduce pain and anxiety caused by treatment [77].

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[11] Clark RA. Fibrin sealant in wound repair: a systematic survey of the literature. Expert Opin Investig Drugs 2000;9:2371 – 92. [12] Cruse PJE, Foord R. A five-year prospective study of 23,649 surgical wounds. Arch Surg 1973;107:206 – 9. [13] Cummings P. Antibiotics to prevent infection in patients with dog bite wounds: a meta-analysis of randomized trials. Ann Emerg Med 1994;23:535 – 40. [14] Demling RH, DeSanti L. Management of partial thickness facial burns. Burns 1999;25:256 – 61. [15] Dire DJ, Welsh AP. A comparison of wound irrigation solutions used in the emergency department. Ann Emerg Med 1990;19:704 – 8. [16] Edlich RF. Tissue adhesives. Ann Emerg Med 1998; 32:274 – 5. [17] Edlich RF. Tissues adhesives: revisited. Ann Emerg Med 1998;31:106 – 7. [18] Edlich RF, Thul J, Prusak M, et al. Studies in the management of the contaminated wound. VII. Am J Surg 1971;117:394 – 7. [19] Falanga V, Margolis D, Alvarez O, et al. Rapid healing of venous stasis ulcers and lack of clinical rejection with an allogenic cultured human skin equivalent: human skin equivalent investigators group. Arch Dermatol 1998;134:293 – 300. [20] Forrest RD. Early history of wound treatment. J R Soc Med 1982;75:198 – 205. [21] Gaboriau HP, Belafsky PC, Pahlavan N, et al. Closure of mucosal defects over exposed mandibular plates using fibrin glue. Arch Facial Plast Surg 1999;1:191 – 4. [22] Galiano RD, Zhao LL, Clemmons DR, et al. Interaction between the insulin-like growth factor family and the integrin receptor family in tissue repair processes: evidence in a rabbit ear dermal ulcer model. J Clin Invest 1996;98:2462 – 8. [23] Gross CW, Gallagher R, Schlosser RJ, et al. Autologous fibrin sealant reduces pain after tonsillectomy. Laryngoscope 2001;111:259 – 63. [24] Greene D, Koch RJ, Goode RL. Efficacy of octyl-2cyanoacrylate tissue glue in blepharoplasty: a prospective controlled study of wound-healing characteristics. Arch Facial Plast Surg 1999;1:292 – 6. [25] Haury B, Rodeheaver G, Vensko J, et al. Debridement: an essential component of traumatic wound care. Am J Surg 1978;135:238 – 42. [26] Hinman CD, Maibach H. Effect of air exposure and occlusion on experimental human skin wounds. Nature 1963;200:377 – 8. [27] Hollander JE, Giarrusso E, Cassara G, et al. Comparison of staples and sutures for closure of scalp lacerations. Acad Emerg Med 1997;4:460 – 1. [28] Hollander JE, Richman PB, Werblud M, et al. Irrigation in facial and scalp lacerations: does it alter outcome? Ann Emerg Med 1988;31:73 – 7. [29] Hollander JE, Singer AJ. Laceration management. Ann Emerg Med 1999;34:356 – 67. [30] Hollander JE, Singer AJ, Valentine S, et al. Wound registry: development and validation. Ann Emerg Med 1995;25:675 – 85.

R. Gassner / Oral Maxillofacial Surg Clin N Am 14 (2002) 95–104 [31] Howard JM, Barker WF, Culbertson WR, et al. Postoperative wound infections: the influence of ultra-violet radiation of the operating rooms and various other factors. Ann Surg 1964;160:32 – 81. [32] Howell JM, Newsome J, Bresnahan K. Histologic effect of butyl-2-cyanoacrylate on skin lacerations. Acad Emerg Med 1996;3:426 – 7. [33] Hudson JW. Wound healing. Oral and Maxifollacial Surgery Clinics of North America 1996;8:457 – 599. [34] Kanegaye JT, Vance CW, Chan L, et al. Comparison of skin stapling devices and standard sutures for pediatric scalp lacerations: a randomized study of cost and time benefits. J Pediatr 1997;130:808 – 13. [35] Krause RS, Moscatti R, Filice M, et al. The effect of injection speed on the pain of lidocaine infiltration. Acad Emerg Med 1997;4:1032 – 5. [36] Kung A. Evaluation of the undesirable side-effects of the surgical use of histoacryl glue with special regard to possible carcinogenicity. Basel, Switzerland: RCC Institute for Contract Research in Toxicology and Ecology, Project 064315; 1986. [37] Langer R. New methods of drug delivery. Science 1990;28:232 – 6. [38] Macsai X, Marian S. The management of corneal trauma: advances in the past twenty-five years. Cornea 2000;19:617 – 24. [39] Majno G. The healing hand: man and wound in the ancient world. Cambridge, MA: Harvard University Press; 1975. [40] Man D, Plosker H, Winland-Brown JE. The use of autologous platelet-rich plasma (platelet gel) and autologous platelet-poor plasma (fibrin glue) in cosmetic surgery. Plast Reconst Surg 2001;107:229 – 37. [41] Manstein CH. What’s wrong with Dermabond? Plast Reconst Surg 1999;104:1587 – 8. [42] Markovchick V. Suture materials and mechanical aftercare. Emerg Med Clin North Am 1992;10:673 – 88. [43] Maw JL, Quinn JV, Wells GA, et al. A prospective comparison of octylcyanoacrylate tissue adhesive and sutures for the closure of head and neck incisions. J Otolaryngol 1997;26:26 – 30. [44] McCaig LF. National hospital ambulatory medical care survey: 1992 emergency department summary. Vital Health Stat 1994;245:1 – 12. [45] Moscati R, Mayrose J, Fincher L, et al. Comparison of normal saline with tap water for wound irrigation. Am J Emerg Med 1998;16:379 – 81. [46] Mustoe TA, Han H. The effect of new technologies on plastic surgery. Arch Surg 1999;134:1178 – 83. [47] Noordiz JP, Foresman PA, Rodeheaver GT, et al. Tissue adhesive wound repair revisited. J Emerg Med 1994;12:645 – 9. [48] Osmond MH, Klassen TP, Quinn JV. Economic comparison of a tissue adhesive and suturing in the repair of pediatric facial lacerations. J Pediatr 1995;126: 892 – 5. [49] Pirisi A. Cosmetically, tissue adhesives are as good as sutures for closing wounds. Lancet 1998;352:1834. [50] Pulapura S, Kohn J. Trends in the development of

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Oral Maxillofacial Surg Clin N Am 14 (2002) 105 – 116

Cartilage regeneration Barbara D. Boyan, PhDa,b,c,*, David D. Dean, PhDa, Christoph H. Lohmann, MDa,d, Gabriele G. Niederauer, PhDe, Jacquelyn McMillan, MBChB, FRCSEd (Orth)a, Victor L. Sylvia, PhDa, Zvi Schwartz, DMD, PhDa,c,f a

Department of Orthopaedics, University of Texas Health Science Center at San Antonio, 7703 Floyd Curl Drive, San Antonio, TX 78229 – 3900, USA b Department of Biochemistry, University of Texas Health Science Center at San Antonio, 7703 Floyd Curl Drive, San Antonio, TX 78229 – 3900, USA c Department of Periodontics, University of Texas Health Science Center at San Antonio, 7703 Floyd Curl Drive, San Antonio, TX 78229 – 3900, USA d Department of Orthopaedics, University of Hamburg Eppendorf, D-20246 Hamburg, Germany e OsteoBiologics, Incorporated, 12500 Network, Suite 112, San Antonio, TX 78249, USA f Department of Periodontics, Hebrew University Hadassah, Jerusalem, Israel

Articular cartilage has been a primary focus of the emerging tissue engineering industry. The market for cartilage repair technologies is considerable for trauma and sports injuries and even greater for regeneration of cartilage in patients with arthritis. Articular cartilage degeneration is accompanied by morbidity, which leads to absenteeism and development of conditions associated with chronic pain. In severe cases, the loss of function is of sufficient magnitude that it is necessary to replace damaged joint tissue with a bioprosthesis.

Problems associated with tissue engineering of cartilage Despite these economic drivers, successful tissue engineering of articular cartilage has been elusive. Unfortunately, cartilage does not heal in the same manner as seen in other tissues, in part because it has only a rudimentary blood supply. When the cartilage is severed, the chondrocytes seal the exposed edges

* Corresponding author. E-mail address: [email protected] (B.D. Boyan).

of the wound and in effect create a new cartilage surface [2,66,72]. The two sides of the defect do not fuse, which creates a focal change in the way that the tissue experiences compressive loads. Similarly, when injuries occur that cause loss of a piece of cartilage, the chondrocytes in the surrounding tissue again seal off the edges of the defect site. There is a limited attempt at repair, but the tissue that forms within the defect tends to be fibrocartilage. There are several reasons why fibrocartilage forms within chondral defects. The source of cells is believed to be synoviocytes [39], which are fibroblastic and as such synthesize type I collagen. Even if chondroprogenitor cells migrate into the defect site, they must produce large amounts of matrix quickly to facilitate migration across large regions of space relative to the cell, and type I collagen is favored under such circumstances. Should the defect penetrate the subchondral plate, a clot is able to form within the defect site and serves as a scaffold for cell attachment and migration. Many of the cells that colonize such defects are derived from the vasculature and marrow stroma [62], however, and possess the capacity to differentiate into various mesenchymal cell types. In the absence of a sufficient supply of chondrogenic factors, these cells tend to select the default pathway and differentiate along a fibroblastic lineage [38,43,49,53,61].

1042-3699/02/$ – see front matter D 2002, Elsevier Science (USA). All rights reserved. PII: S 1 0 4 2 - 3 6 9 9 ( 0 2 ) 0 0 0 1 7 - 1

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Articular cartilage architecture In comparison with organs such as liver and kidney, cartilage is a deceptively simple tissue in that only one cell type is present. This is both strength and weakness with respect to tissue engineering. Although there is no need to design complex three-dimensional constructs that contain several different types of cells, the heterogeneity within the cartilage in terms of chondrocyte phenotype does require consideration. Cells at the top of the articular cartilage exhibit a flattened morphology. These cells form the barrier with the synovial fluid and provide a surface that can withstand shear. Most of the cartilage is hyaline in nature, with clonal populations of chondrocytes, each in its own lacunae, surrounded by a pericellular matrix that differs from the interterritorial matrix. The cells adjacent to the subchondral bone have yet another phenotype. These cells are hypertrophic and appear in many ways like a miniature growth plate. This region is called the tidemark because it is characterized by a distinctly different staining property, partly because of the difference in the extracellular matrix and the deposition of hydroxyapatite. This phenotypic heterogeneity of the chondrocytes is accompanied by a complex three-dimensional architecture provided by the extracellular matrix. Type II collagen fibrils delineate the form of the matrix, which results in overlapping gothic cathedral-like arches that separate columns of cells [19,40]. There is also secondary structure in the form of crossstruts. In addition to the collagen framework, there is a fine net formed of cartilage oligomeric matrix protein (COMP) throughout the interterritorial matrix [26,75]. Interspersed among the fibrillar proteins are proteoglycan aggregates, composed of a hyaluronic acid backbone to which the aggrecan monomer is connected via association with link protein [9]. The number of aggrecan monomers and the length of the hyaluronic acid chain vary with the articular cartilage zone. Aggrecan is decorated with sulfated glycosaminoglycan side chains that also vary in number and length. During embryonic bone formation, the articular cartilages form as remnants of cartilage anlagen as they are replaced with bone and marrow. The embryonic tissue is highly cellular, but postfetal cartilage is relatively acellular. The articular cartilage continues to grow as the long bones continue to lengthen; however, most cartilage repair is completed in adults, and by then, any mitosis in primary cartilages is infrequent. It is likely that chondrocytes retain some proliferative capacity, which is stimulated by injury. Even so, the neocartilage that forms within a chondral

defect has an extracellular matrix that is different from that of the cartilage surrounding the defect site. The architectural features of the matrix are honed over time and occur in response to directional loads that are placed on the tissue [67,73]. Even the types of proteoglycan in the extracellular matrix may differ in terms of core protein [18] and glycosaminoglycan sulfation [17,34]. The amount of hydration of the tissue and, consequently, the hydrostatic forces experienced by the cells differ [35,46]. The resistance of the tissue to compressive loads also differs [5]. The integration of the neocartilage with the surrounding cartilage is often ineffective, not only because of the differences in matrix composition but also because of the cartilage matrix sealing effect. This results in mechanical instability that further reduces the success of the repair tissue. In healthy joints, the base of the articular cartilage is ‘‘glued’’ to the subchondral bone in a manner that is similar to the way that the base of a growth plate is attached to the bony metaphysis. By calcifying their matrix, the chondrocytes in the tidemark region form an interlock with the mineralized bone. Disruption of this interface also contributes to mechanical instability. It is currently known that it is critical to have good repair of the subchondral plate to have good repair of the articular cartilage [44]. Unfortunately, the subchondral plate can be damaged through trauma, during surgery, or as a consequence of degenerative diseases. Tissue-engineering strategies must consider how to ensure that mineralization of the cartilage occurs in an appropriate manner and provide a method for regenerating the bone itself.

Cartilage calcification and the tidemark Physiologic mineralization of cartilage involves chondrocyte hypertrophy and deposition of apatite crystals within the extracellular matrix. Both processes require turnover of the extracellular matrix. As the cells produce matrix, they also synthesize matrix processing enzymes and store them in inactive form as zymogens or together with inhibitors such as tissue inhibitor of metalloproteinase (TIMP) [14]. Other matrix processing enzymes are packaged in extracellular matrix vesicles and are released in active form upon signals from the cells. Because of the importance of maintaining a functional architecture, matrix turnover is slow in healthy tissue. At the tidemark, however, the rate of matrix degradation is believed to be increased, both to accommodate hypertrophy and to redesign the matrix for mineral deposition [21 – 23]. These cells produce increased numbers of matrix

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vesicles, which also provide sites for the initial formation of calcium phosphate crystals [55]. Successful tissue engineering must result in a cartilage in which these events occur in a regulated manner in time and space. Should the hypertrophy of the cells extend too far up into the hyaline cartilage, the possibility of osteophyte formation is increased, as is the likelihood of tissue destruction. If hypertrophy and matrix mineralization do not occur, then the repair cartilage is likely to delaminate.

Tissue engineering of cartilage This section of the article focuses on the engineering of articular cartilages at the ends of long bones. The mandibular condyle is a secondary cartilage and as such retains the ability to remodel throughout life. Relatively little is known about the cellular, biochemical, and structural features of the articular cartilage of the condyle that impinge on the success of tissue engineering in this tissue. For this reason, engineering strategies for the mandibular cartilage are not discussed. The need for cells The repair capacity for cartilage is low, because of the relatively low proliferative capacity of cells resident in the tissue and their low migratory ability. Because of this, the strategies for cartilage tissue engineering have focused on the delivery of cells to the defect site through various approaches. There are three different types of cartilage defects, each of which requires its own set of repair strategies. Partial-thickness defects Defects that do not extend to the subchondral plate are called ‘‘chondral’’ or ‘‘partial-thickness’’ defects. Surgeons are reluctant to use strategies that penetrate the subchondral plate for repair of these types of defects because to do so requires loss of additional cartilage. The usual approach is to remove undermined chondral flaps from the edges of the defect and leave stable cartilage undisturbed and hope for the best. In some instances, this is not a practical choice, yet the reluctance remains. The problems of how to maintain cells at the site or ‘‘glue’’ allograft to the existing cartilage have not yet been resolved. Hunziker and Rosenberg [39] have characterized the healing process of partial-thickness defects with respect to the cells that colonize the defect site and methods for enhancing their ability to differentiate into chondroblasts and ultimately become functional

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chondrocytes. Preparation of the chondral defect is critical. Not only is it necessary to remove the damaged cartilage but also healing is enhanced if the walls of the defect are etched with enzymes that remove the outer layer of cartilage matrix proteins. Host cells adhere more effectively to the tissue surface. They also have noted that delivery of transforming growth factor-beta (TGF-b) to the site in various time-release carriers also improves the differentiation of the colonizing cells into chondrocytes. Full-thickness defects When the cartilage defect extends to the subchondral plate, several strategies can be considered. The choice of strategies is related to the actual size of the defect. In humans, trauma and sports injuries result in defects that vary in size but tend to be focal. In contrast, arthritis leads to large regions of degenerate tissue and has the added complication of being chronic, involving the subchondral bone, and compromising the surrounding ‘‘healthy’’ cartilage tissue. The biggest problems to be resolved in treating full-thickness defects are the interface with the surrounding cartilage and the tendency of the neocartilage to delaminate. Microfracture. By far the most common tissue engineering strategy relies on microfracture of the subchondral plate, providing access of cells from the underlying vasculature and marrow stroma to migrate into the defect site. This technique has been popularized by Blevins et al. [8] and Steadman et al. [65] with reasonable success. After removing all damaged cartilage, the subchondral plate is penetrated at several sites. The blood that moves into the synovial space forms a clot in the defect and brings with it growth factors that can enhance tissue repair. Mesenchymal cells migrate onto and through the hematoma, and in the presence of the surrounding cartilage at least some of these cells differentiate into chondroblasts. The subchondral plate is repaired through endochondral ossification, the same mechanism that repairs bone fractures. The chondroblasts produce a cartilage matrix, but the quality of the matrix is variable at best. Eventually, fibrocartilage predominates. This approach can be successful clinically because the patient can return to function relatively pain free. Over time, the repair tissue may fail, but for some patients, it is adequate. Periosteum and perichondrium. The periosteum, the skin-like tissue that forms on the outer surface of bone, is a rich source of multipotent mesenchymal cells, including cells with the potential to become chondrocytes. O’Driscoll et al. [50 – 52] have pio-

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neered the use of periosteal autografts for treatment of chondral defects. By placing the periosteum in an inverted fashion in the defect site, the mesenchymal cells can migrate easily into the defect space. Their ability to migrate out of the site is reduced by the graft itself. Ideally, the periosteal cells differentiate into chondroblasts in the presence of chondrogenic signals from the surrounding cartilage. This technique has proved successful in some cases, but exactly what makes one case a success and another case a failure is not well understood. Pretreatment of the tissue with TGF-b improves chondrogenesis [45] for the same reasons as noted for partial-thickness defects. To improve on the concept, Amiel et al. [1] and Chu et al. [20] have suggested that perichondrial tissue may be a better source of cells. The perichondrium is an extension of the periosteum, but it overlies cartilage tissue. These investigators reasoned that it was likely to contain a greater percentage of chondroprogenitor cells. They have used perichondrium successfully as a source of cells for treatment of full-thickness defects. Again, treatment of the cells with TGF-b improved success. Cartilage cell therapy Another approach is to use committed chondrocytes rather than tissue that contains chondroprogenitor cells. Allografts and ex vivo tissue-engineered cartilage Cartilage autografts are in short supply; otherwise the patient would not need a tissue-engineering approach for repair of the defect. Allografts are used to cover large defects, with varying success [6,16]. One method that holds promise is the use of allografts produced by tissue engineering ex vivo. Several methods have been developed to achieve the goal of making a tissue in culture that has the properties of articular cartilage and integrates with the surrounding host cartilage and with the subchondral bone [68,71]. To accomplish this goal, several problems have emerged that have required innovative solutions, including the need for a reliable source of cells, the need for a structural scaffold on which to grow the cells in three dimensions, the contribution of mechanical force to the development of the tissue, and the development of novel bioreactors to achieve sufficient tissue in a reasonable period of time [36,56]. Despite considerable success in solving each of these problems, the resulting tissue still does not replicate the qualities of the native host tissue. The only Food and Drug Administration-approved cartilage cell therapy currently in use clinically in the

United States is the direct delivery of autologous chondrocytes to the defect site. This methodology was pioneered by Brittberg et al. [15] and has been the subject of several studies over the past 5 years. The concept is simple. Cartilage is removed surgically from a non – weight-bearing region of the joint to be treated, and the cells isolated from that cartilage are cultured ex vivo through four passages. This practice is critical because of the low cellularity of cartilage. It is necessary to expand the number of cells to increase the likelihood of repopulating the defect with chondrocytes. At the same time of this surgery, the defect site is prepared. Periosteal tissue is removed from the bone and used to create a seal over the surface of the defect. Currently, this periosteal flap is sutured, but the process of suturing itself leads to defects in the cartilage, so attempts to develop a glue to attach the periosteum to the cartilage surface are also underway. While the cells are in culture, the surgical site is allowed to heal. When the cells are ready, they are injected into the defect site arthroscopically. Technically, the periosteal seal is intended to retain the cells within the defect. Practically, this does not occur. Second-look arthroscopies show that the periosteum can be dislodged and that the cells tend to migrate out of the defect. When the treatment does work, however, the results are promising. Whether they hold up in the long term is not yet known. Studies using dogs and goats suggest that this neocartilage brings with it the same problems as seen with other repair strategies [13,24]. It also is not clear whether the repair cartilage is the consequence of the cells in the periosteum, the cells that are injected, or some combination of the two. More recently, investigators have begun to grow cartilage in multilayers on membranes [41] and to transfer the construct to the defect site rather than to inject cell suspensions. This technique may not require the use of a periosteal seal, because the cells are in effect anchored to the underlying membrane. The limitation is the attachment to the subchondral bone. All of the cell strategies are limited by the tendency of chondrocytes to lose their chondrogenic phenotype in monolayer culture. More than 20 years ago, Benya et al. [7] noted that articular chondrocytes lose their ability to synthesize type II collagen and produce sulfated proteoglycan when they were grown in monolayer culture. The chondrocyte phenotype was restored when the cells were grown in suspension culture or in various kinds of gels. Alginate was particularly effective at preserving these cell characteristics [31]. The reasons for this are not entirely understood, but it is clear that retention of the rounded articular chondrocyte morphology is impor-

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tant. Unfortunately, cells grow slowly using these methods, which makes it difficult to achieve sufficient expansion. Recently, alginate has been used to microencapsulate articular chondrocytes after culture in monolayer and found that the microencapsulation process restored the ability of the cells to produce type II collagen and sulfated proteoglycan (A. Sambanis, unpublished observation). This opens up the possibility of expanding the cells in monolayer culture and then preparing them for injection via microencapsulation. Alternatively, articular chondrocytes can be expanded in bioreactors. Various methods are being developed for this purpose. Currently, the cells tend to form spherical nodules in these sorts of culture systems and develop three-dimensional architecture within the nodules, including hypertrophic chondrocytes at the center [28,69,74]. Whether these nodules are suitable for use in treating full-thickness defects is not yet clear. Another approach is to use chondrocytes that do not undergo such marked loss of phenotype in vitro. The authors’ laboratory has used hyaline chondrocytes from the resting zone of the costochondral cartilage of the rib for this purpose. Although these cells are not identical to articular chondrocytes, they are not yet in the endochondral lineage per se, and they produce a proteoglycan-rich extracellular matrix through four passages in culture [10]. Although they do produce

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type I collagen as an adaptation to monolayer culture, they also continue to produce type II collagen [10], and when they are implanted in nude mouse muscle in vivo, they form cartilage nodules that are indistinguishable from normal hyaline cartilage [11]. One problem with using chondrocytes is their tendency to undergo hypertrophy (Fig. 1). To increase the number of chondrocytes while at the same time control their differentiation along the endochondral pathway is an elusive goal of cartilage cell therapy. One approach is to pretreat autologous chondrocytes with growth factors that modulate one or more of the phenotypic characteristics of the cells. Unfortunately, growth factors are pleiotropic, affecting different cells in different ways. They may affect the same cell differentially depending on the cell’s state of maturation within its lineage cascade. This is the case with chondrocytes. As shown in Table 1, cells from the resting zone of costochondral cartilage respond to various regulatory factors in a manner that is distinct from the response of growth zone cells, which are derived from the prehypertrophic and upper hypertrophic zones of the rat costochondral growth plate. Table 1 is a summary of several experiments performed in the authors’ laboratory using factors and hormones that are known to regulate cartilage and bone or that are being tested for their effectiveness in various musculoskeletal tissue-engineering applications [42,43,58 – 60]. Many of the factors increase

Fig. 1. Rat costochondral cartilage resting zone chondrocytes were loaded on a polylactic acid/polyglycolic acid porous foam scaffold and implanted for 8 weeks in the calf muscle of a nude mouse. Chondrogenesis has occurred throughout the scaffold space and where the scaffold has resorbed, the cartilage is in direct contact with the surrounding muscle. Chondrocytes at the center of the neocartilage are producing a cartilage matrix and some of the cells are undergoing hypertrophy. Residual scaffold is evident as a clear area, due to its dissolution during processing. (Haematoxylin and eosin, original magnification
10).

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Table 1 Cell maturation and chondrocyte response to bioactive factors EMD

TP508

PDGF-BB

IGF-1

TGF-b1

FGF-2

BMP-2

1,25

24,25

Resting zone cells Proliferation Alkaline phosphatase [35S]-sulfate incorporation Response to 1,25-D3

" # — ND

— # " ND

" — " —

" " " —

" " " "

" " " ND

" — " "

# — — ND

"# " " "

Growth zone cells Proliferation Alkaline phosphatase [35S]-sulfate incorporation

" — #

" — —

ND ND ND

ND ND ND

" " "

ND ND ND

" " —

# " "

" — —

Abbreviations: FGF-2, basic fibroblast growth factor; IGF-1, insulin-like growth factor; ND, not done; 1,25, 1,25 dihydroxyvitamin D3; 24,25, 24,25 dihydroxyvitamin D3.

proliferation of the hyaline-like resting zone cells. Of those factors that stimulate proliferation, some also stimulate proteoglycan production based on the incorporation of radiolabeled sulfate into glycosaminoglycans, which indicates that cartilage matrix synthesis is enhanced. Most of the factors also promote endochondral differentiation of the cells, however. Alkaline phosphatase specific activity is increased, and this enzyme, which is associated with extracellular organelles called matrix vesicles, is involved in calcification. Resting zone cells normally respond primarily to the vitamin D metabolite 24R,25(OH)2D3, whereas the more mature growth zone cells respond primarily to 1a,25(OH)2D3. When resting zone cells are treated with TGF-b1, bone morphogenetic protein2, or 24R,25(OH)2D3 for extended periods of time, however, responsiveness to1a,25(OH)2D3 is upregulated and responsiveness to 24R,25(OH)2D3 is lost, which indicates that these cells not only have become more differentiated but also have acquired a growth zone chondrocyte phenotype [58 – 60]. The factors that show the most promise for pretreatment of chondrocytes are factors that stimulate proliferation and cartilage matrix production but retard or inhibit endochondral differentiation. Of all of the factors that the authors have tested to date, the best candidates seem to be Emdogain (EMD) (Biora, Inc., Malmo, Sweden), which increases the pool of chondroprogenitor cells [12,57], TP508 (Chrysalin, Chrysalis, BioTechnology, Inc, Galveston, TX), which enhances matrix synthesis but not differentiation, and platelet-derived growth factor-BB (PDGF-BB), which stimulates proliferation and matrix synthesis, but not endochondral maturation [42]. To test whether these in vitro assays are relevant to in vivo behavior, the authors pretreated resting zone chondrocytes for 4 hours or 24 hours with PDGF-BB

before implantation in nude mouse muscle. For these experiments, they used scaffolds provided by OsteoBiologics, Inc. (San Antonio, TX). After treatment, the cells were loaded onto the scaffolds and implanted bilaterally in the calf muscles. Implanted tissue was examined histomorphometrically for the presence of new cartilage and the degree of chondrocyte hypertrophy at 4 and 8 weeks. The results showed that PDGF-BB pretreatment increased the amount of neocartilage at 8 weeks and prevented chondrocyte hypertrophy; however, the shorter exposure to the growth factor was more effective than the longer exposure [43]. This result shows that the type of factor used is important, and it indicates that the pretreatment regimen may be critical. Finally, several other cell-based strategies are in various states of development. Mesenchymal stem cells (MSCs) were first used by Wakitani et al. [70] to repair full-thickness defects in rabbits. These multipotent cells differentiated into chondrocytes and formed neocartilage within the experimental defects that was comparable to neocartilage formed by committed chondrocytes. The effectiveness of the strategy was improved through the use of gels to create MSCbased constructs [54,64]. The use of MSCs is a powerful technology because of their high proliferative capacity and their ability to differentiate into tissues. There are some caveats, however. For example, recent studies suggest that the phenotype of MSCs can vary, depending on the physiology of the donor [47]. Even taking this into consideration, MSCs hold great promise. One of the most innovative uses is the use of fat as a source of MSCs [25,76]. MSCs in fat possess the ability to differentiate into chondrocytes when treated with growth factors that induced chondrogenesis, such as TGF-b. Similarly, MSCs isolated from perichondrium can be used if pretreated

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Fig. 2. Mosaic arthroplasty showing the placement of osteochondral cores within a defect. (Courtesy of Smith and Nephew, Endoscopy Division, Andover, MA.)

with TGF-b or when transfected with plasmids that contain TGF-b1 DNA [32]. Osteochondral defects In instances in which the subchondral bone is damaged, a core of cartilage and underlying bone is removed. If the defect is small enough, healing occurs by a process similar to that described for microfracture. In larger defects, the repair of the subchondral bone is inadequate to support the repair cartilage, and the result is poor. For these types of defects, mosaic arthroplasty is preferred. Mosaic arthroplasty. Mosaic arthroplasty relies on the use of an osteochondral plug from a region of the articular cartilage that is normally not required for weight bearing [36,37]. The surgeon may remove several cores of cartilage and bone from this site and then place them in a mosaic-like fashion in the defect site on the weight-bearing region of the joint (Fig. 2). This method works well because the cartilage in each plug remains viable, but problems arise because the transplanted cartilage does not fuse with the surrounding host tissue. Just as the edges of the cartilage defect site seal, so do the edges of the cartilage plugs. The defect site is osteochondral and permits clot formation and osteochondral repair around and between each plug. This repair tissue is fibrocartilage as would be the case for microfracture. The final result is an improvement over simply allowing the larger osteochondral defect to heal via clot formation and mesenchymal cell defferentiation, but the me-

chanical instabilities created by the variation in tissue type may predispose this strategy to failure. Cell-based therapies. Cell-based therapies also are used for osteochondral repair, but they depend on a suitable delivery device because the subchondral plate is not present to serve as a substrate. All of the same cell strategies as described for full-thickness defects with respect to cell source are applicable. The physical properties of the scaffold are critical to ensure appropriate healing of the defect with bone and cartilage, however. Simple biodegradable felts work well for growing cartilage allografts ex vivo for use in full-thickness defects [27,29,36], because all that is required is a structural support for the cells. In contrast, the scaffolds used in osteochondral defects also must possess mechanical properties. In small defects, a singlephase implant is possible, because the marrow stromal cells that populate the scaffold quickly form new bone and provide a substrate for the overlying neocartilage. In larger defects, however, two-phase implants are preferred, with each phase mimicking the structural and mechanical properties of the tissue with which it interfaces. Various versions of a two-phase implant are in development. Frenkel et al. [30] used type I collagen as the structural base of the implant, varying the density of the collagen fiber mesh, with looser fibers in the cartilage phase and more compacted fibers in the bone phase. At the interface the network of fibers is so dense that it serves as a barrier for migration of

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Fig. 3. Placement of a two-phase polylactic acid/polyglycolic acid osteochondral implant in the condyle. The stiffer bone phase interfaces with the subchondral bone, whereas the cartilage interfaces with a scaffold that mimics its mechanical properties. In this particular implant, the cartilage phase is covered with a surface that is designed to reduce shear stress. (Courtesy of OsteoBiologics, Inc., San Antonio, TX.)

chondrocytes down into the bone phase or for migration of marrow stromal cells up into the carti-

lage phase. This design is intended to mimic the barrier provided by the subchondral bony plate. When articular chondrocytes are precultured on the implant before its placement in an osteochondral defect, healing is excellent in animal models. This design also has been produced as a biodegradable implant constructed from polylactic acid fibers [33]. The implant is anchored in the site with small polylactic acid pins. The authors have designed a two-phase implant that not only varies pore size, as would be achieved by modulating the packing density of collagen fibers, but also varies mechanical properties of each phase [4]. This implant is constructed using polylactic acid and polyglycolic acid to create a foam that is stiffer in the bone phase than it is in the cartilage phase. Studies that use rabbits as the animal model show that the material properties of the neocartilage are comparable to those of the surrounding host cartilage when implants of this design are used, particularly when TGF-b1 is incorporated into the cartilage phase. The importance of retaining the mechanical properties of such an implant over the initial healing period was demonstrated in a study using goats as the animal model [3]. Compression of the scaffold also led to premature compression of the neocartilage and failure of the repair strategy, however. In vitro studies were conducted to show that fiber reinforcement can be used to tailor the mech-

Fig. 4. Typical histological appearance near the center of an osteochondral defect that was treated using a two-phase implant to which chondrocytes had been added prior to implantation. The type of experimental defect used in this study was ‘‘critical sized’’ so complete healing across the entire defect space was not anticipated. Subchondral bone is restored and there is excellent integration of neocartilage with the surrounding host tissue. Defects treated with implants that were not preloaded with cells also had a similar appearance. (Courtesy of OsteoBiologics, Inc., San Antonio, TX.)

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anical properties of porous, resorbable scaffolds for optimal performance in an articular cartilage environment [63]. A study using a goat femur model to test these newly developed synthetic, resorbable multiphase implants designed specifically for osteochondral cartilage repair showed the value of this approach [48]. In this screening study, four different multiphase implants with varying mechanical and physical properties were randomly implanted in weight-bearing (high load environment) (Fig. 3) and non – weight-bearing (low load environment), 3-mm diameter defects (n = 64) of goat distal femurs. The implants were assessed for their effectiveness as cartilage repair scaffolds after 16 weeks of healing using gross, biomechanical, and histologic evaluations. All implants were tested as scaffold alone (cell-free) and loaded with autologous costochondral chondrocytes. Qualitative histologic evaluations showed that all groups had a high percentage of hyaline cartilage and good bony restoration, with new tissue integrating well with the native cartilage (Fig. 4). Defect healing in the condyle was significantly higher than in the patellar groove, but there were no differences in healing because of implant type. In weight-bearing sites, the quality of the neocartilage was equally good regardless of whether cells were added to the scaffolds before implantation. This observation suggests that cells resident in the host cartilage may contribute to the formation of neocartilage at the interface of the scaffold and the native tissue when the mechanical environment is favorable, thereby achieving the goal of integration. Marrow stromal cells also can differentiate into hyaline and hypertrophic chondrocytes in an appropriate manner, again when attaching to a physical substrate that is like that of the natural tissue.

Summary This article has shown the problems and challenges of tissue engineering cartilage and has presented the current strategies that are under investigation. The specific characteristics of the tissue are advantages and disadvantages. Surgeries can be performed arthroscopically, but the lack of a robust intrinsic healing response hampers the effectiveness of the therapy. We have not yet solved the problem of the choice of cells, nor have we identified all of the requisites for optimal scaffold design. Efforts thus far have focused on small defects in relatively healthy patients. How aging, disease, and pharmacologic intervention will modify the effectiveness of tissue-engineered cartilage,

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whether it is produced in vivo or ex vivo, is still unknown. Despite the problems, the advances made over the past 5 years suggest that the challenges will be met.

Acknowledgements The authors thank Sandra Messier for her assistance in the preparation of the manuscript, and they acknowledge the support of the National Institutes of Health (US PHS grants DE-08603 and DE-05937). The authors also thank Dr. Frank Barry, Osiris Therapeutics, Inc. (Baltimore, MD) for his assistance in the preparation of the manuscript; Smith and Nephew (Andover, MA) for their contribution of Fig. 2; and OsteoBiologics, Inc. (San Antonio, TX) for their contribution of Figs. 3 and 4.

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Peripheral sensory nerve regeneration with biodegradable materials and neurotropic factor Arden K. Hegtvedt, DDS, MS a, John R. Zuniga, DMD, MS, PhD b,*, Erick M. Rath, DDS, PhD c a

Department of Oral and Maxillofacial Surgery, The Ohio State University, 305 W. 12th Avenue, Columbus, OH 43218, USA b Department of Oral and Maxillofacial Surgery, University of North Carolina School of Dentistry, Campus Box 7450, Chapel Hill, NC 27599, USA c Black Hills Oral And Maxillofacial Surgery, 3415 Fifth St., Rapid City, SD 57701, USA

Peripheral nerve injuries that meet the criteria of a Class 5 Sunderland’s or Seddon’s neurotemesis include those injuries for which there is discontinuity of the nerve. Every microsurgeon is keenly aware that these injuries may inhibit the ability to approximate proximal and distal nerve stumps during microsurgical repair because of insufficient nerve to create a passive coaptation. Various techniques have been described for bridging an avulsive gap, including autogenous peripheral nerve grafts taken from distant sites (i.e., sural, greater auricular, medial antebrachial cutaneous nerve), which are considered to be standard clinical practice [11,14 – 17,26,42,44,62]. The results of autogenous grafting have not proved optimal [35] and have been accompanied by pain, unacceptable paresthesia, neuroma development, and visible scar formation at the donor site [34,64]. A need exists for alternative nerve grafting techniques to circumvent these problems while providing biologic properties comparable to autologous tissue. Entubulation repair was first advocated for peripheral nerve repair because it was simple and biologic materials became available that met the criteria necessary for implantation in humans. The proposed advantages were that fibroblastic growth into the repair site would be limited, directional guidance to regenerating axons was provided, and trophic factors from the injury site would concentrate within the tube to facilitate regeneration [6,10,18,28,29,31,32,34,36 – 41,43,45,53,60,65]. Early studies involved the silicone

* Corresponding author.

prostheses [28,29,32,54,65], but it was not resorbable and prevented exchange of biologic materials into or out of the tubes, which resulted in decreased regeneration of axons. Collagen tubes were developed because they were semipermeable and bioresorbable [4,53,60]; however, they had limited permeability and allowed for connective tissue ingrowth and tube collapse. Biodegradable polyglycolic acid mesh grafts were introduced and offered larger pore size and flexibility for easy placement [10,18,34,40,41]. When applied to human digital nerve repairs, the polyglycolic entubulation method was as successful as autologous nerve grafts; however, entubulated grafts larger than 3 cm were inferior to autogenous nerve grafts [34]. Polyglycolic entubulation in a 1.2-cm inferior alveolar nerve defect in a 51-year-old man failed to provide functional sensation [5]. Gore-Tex (Gore Company, Flagstaff, AZ) tubes used to reconstruct seven inferior aveolar and lingual nerve injuries in humans were not effective except in defects less than 3 mm in length, in which the material acted as protective barrier membrane rather than as a conduit [47]. Entubulation with biodegradable materials seems to offer a reasonable alternative to repair with autologous nerve tissue. Disadvantages of the available systems include evidence of the failure of sufficient volumes of proximal nerve fibers to bridge to the distal nerve stump for adequate sensory recovery. The authors hypothesize that this is partially because of the passive conduit role of these prostheses. The development of biocompatible materials, which actively enhance (neurotrophic) or support (neurotropic) nerve growth in conjunction with the conduit provided by

1042-3699/02/$ – see front matter D 2002, Elsevier Science (USA). All rights reserved. PII: S 1 0 4 2 - 3 6 9 9 ( 0 2 ) 0 0 0 1 4 - 6

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Fig. 1. Sensory testing apparatus. The tip of the thermode is applied to the rabbit’s lower lip. A stop watch is used to determine the time elapsed from application of the stimulus to response. The control box contains a rheostat for adjustment of temperature. A thermistor runs from the tip of the thermode to the control box, yielding a temperature display on the box to provide temperature feedback at the tip.

entubulation, potentially will increase the success in clinical situations. Longer nerve gap distances (20 – 25 mm) in the rat sciatic nerve have been bridged using silicone tubular prostheses when Matrigel (Collaborative Research Inc, Bedford, MA) was placed into the lumen [63]. Matrigel is biosynthetic basement membrane material derived from a hybridoma. It contains laminin, collagen type IV collagen, heparin sulfate, proteoglycan, and enactin. Laminin, one of the principal components of Matrigel, is solubilized basement membrane that has been shown to act as substrate for nerve growth [12,23]. Nerve growth is supported by enhanced attachment and differentiation of normal and transformed anchorage-dependent epitheloid cell

types. The purpose of this study was to examine the effectiveness of autogenous and alloplastic grafts as means of supporting nerve regeneration of the rabbit inferior alveolar nerve. The hypothesis that drove this study was that polyglycolic mesh entubulation and Matrigel both support nerve regeneration and the combined use provides effective nerve regeneration and sensory recovery when compared to autogenous nerve grafting.

Materials and method Animals Twenty-six adult female New Zealand White rabbits were obtained from the Franklington Rabbitry,

Fig. 2. The course of the inferior alveolar/mental nerve in a rabbit is shown. The dashed rectangle indicates the region of the decortation of the lateral mandibular cortex for exposure of the inferior alveolar nerve (IAN).

Fig. 3. The placement of a graft between the proximal and distal nerve stumps using microsurgical technique.

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Fig. 4. Vicryl mesh grafts are prepared from a sheet of Polyglactin 910 Mesh (Johnson & Johnson, Ethicon, Sommerville, NJ).

placed in individual cages in the University of North Carolina animal quarters, and fed rabbit chow ad libitum. The animals were kept on a 12:12 hour light: dark cycle. Five rabbits were sacrificed before completion of the experimental protocol for various reasons and were eliminated from the study, leaving 21 rabbits for study. The rabbits were divided into five groups: (1) sham (4 rabbits, eight inferior alveolar nerve (IAN) were decorticated from the lateral mandible), (2) control (4 rabbits, eight Inferior alveolar nerve (IAN) were decorticated from the lateral mandible and 1.1-cm segment of the nerve was removed and not replaced), (3) autogenous (4 rabbits, eight IAN were decorticated from the lateral mandible and 1.1-cm segment of nerve was removed, inverted, and repaired to gap the surgical defect), (4) vicryl

mesh (4 rabbits, eight IAN were decorticated from the lateral mandible and a 1.1-cm segment of nerve was removed and the defect closed with polyglactin 910 vicryl mesh (Johnson & Johnson Ethicon, Sommerville, NJ)), and (5) vicryl mesh + Matrigel (5 rabbbits, ten IAN decorticated from the lateral mandible and a 1.1-cm segment of nerve removed and the defect closed with vicryl mesh with 0.1 mL of Matrigel). Presurgical sensory training Before surgery, the rabbits were tested for thermal sensation over the IAN distribution by application of the tip of a thermode to the lower lip at the junction of the vermillion and the mucosa (Fig. 1). The thermode was designed to deliver escalating temperatures to a tip

Fig. 5. The size of the graft is adjusted so that it is slightly longer than the nerve gap to afford overlap of the proximal and distal nerve stumps within the mesh.

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Fig. 6. The vicryl mesh is sutured to the ends of the nerve using an epineurial suture to secure the graft to each nerve stump. An additional suture is used to sew the graft to itself to form a tubular structure.

while providing feedback to ensure precise control of the temperature at the tip of the thermode. Sensory testing was performed while the rabbit’s head protruded from a rabbit-restraining device without the need for anesthesia. The activated thermode delivered an initial temperature of 40°C and rose at a rate of 1°C per second up to 80°C. The thermode was activated once skin contact with the tip was made. Two responses were recorded: (1) Thermal pain threshold was the temperature at which the animal first responded adversively by grimace, withdrawal, or licking. (2) Thermal pain tolerance time was the time it took the animal to respond to a temperature 6° above

threshold applied for 10 seconds. Five semi-weekly testing sessions were compiled for each rabbit, each group, and the entire study population before surgery to ensure that each rabbit’s baseline threshold was recorded as the temperature at which the rabbit’s response was duplicated in at least two consecutive sessions before surgery. Surgery All surgical procedures were performed under general anesthesia using xylazine 10 mg/kg, acepromazine 1 mg/kg, and ketamine 40mg/kg intra-

Fig. 7. The graft is secured to both nerve stumps so that the lumen of the prosthesis is exposed. Matrigel is then applied to the lumen of the vicryl mesh prosthesis by injection with a 30 gauge needle on a 1 mL syringe. The Matrigel is allowed to thicken prior to application because a low viscosity is necessary to provide adequate injection and retention within the tubular mesh.

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Fig. 8. The tube is then completely formed by placing interrupted sutures across the edges of the vicryl mesh.

muscularly in the gluteus muscle. Supplemental intramuscular anesthesia was provided by giving an additonal one half of the induction dose when anesthesia was determined to be light. The surgical exposure of the mandible was performed by a submandibular exposure of the lateral mandible. The mental nerve was exposed, and a combination of rotary instruments and chisels was used to expose the inferior alveolar nerve within the canal (Fig. 2). The inferior alveolar nerve was randomly assigned to one of five groups according to a random block design. All nerve repairs and entubulation procedures were performed with 10-0 monofilament nylon suture (Ethicon, Sommerville, NJ) on a BV1305 tapered needle using a surgical microscope (Fig. 3).

In the autogenous group, the 1.1-cm segment of IAN was removed, rotated distal to proximal (180° on itself), and sutured. When vicryl mesh was used for nerve repair, it was cut from a large sheet of woven polyglactin 910 material and formed into a tubular prosthesis spanning the gap and overlapping the ends of the epinerium (Figs. 4 – 8). When Matrigel was used, it was applied to the lumen of the tubular prosthesis by injection. Matrigel was stored in a glass bottle and kept frozen until use. Upon thawing, the Matrigel enters a liquid state and then gradually increases in viscosity until it reaches a firm gel consistency (similar to agar) so that the placement and retention into the tubular prosthesis were assured.

Fig. 9. The regions from which sections of nerve tissue are removed for histologic examination are shown as the proximal, middle, and distal.

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Fig. 10. The temperature at which the animal first responds to the thermal stimulus is plotted for the five presurgical sensory testing sessions described in the text. This data indicates that the rabbits are learning to respond when exponsed to a painful stimulus. Surgery is delayed until the rabbit demonstrates repeatable thermal threshold values for at least two consecutive tests.

Postsurgical sensory testing The postsurgical animals underwent thermal pain threshold and tolerance time testing twice weekly for 9 weeks. Thermal percent maximum possible effect (T%MPE) was defined as the time from application of stimulus to response compared to baseline using the formula T%MPE = (TL-BL)/ML-BL)  100, where TL was test latency, BL was basal latency, and ML was maximal latency (or 10 seconds). The animals were sacrificed 10 weeks after surgery and perfused with 2% paraformaldehyde/0.5% glutaraldehyde (pH 7.4). The nerves were removed carefully intact and stored in perfusion solution. Histologic preparations were completed on crosssections of each nerve at three sites: (1) 2 to 3 mm proximal to the exposed, injured, or grafted section, (2) in the middle, and (3) 2 to 3 mm distal to the same area (Fig. 9). Fifty-micron sections were taken using a vibratome or cryostat. Sections were processed for light and electron microscopic examination by incubating sections in 2% osmium tetroxide, dehydrated in ascending concentrations of alcohol, and plastic embedded with Medcast. One-micron sections were stained for myelin with methylene blue-azure II, and myelinated fibers were examined using light micro-

scopy. One hundred-nanometer sections were poststained with 5% uranyl acetate and lead citrate and examined using a transmission microscope for unmyelinated fiber content. Morphometric data were generated from photomicrographs using a computer-assisted technique previously described by Zuniga [67]. Myelinated fiber perimeter, axon perimeter, g-ratio (myelin thickness), myelinated fiber cross-sectional area, axon cross-sectional area, myelin/axon ratio, numerical density, surface density, and volume density were determined for each section in each group. Preliminary surface and numerical density data were subjected to Nested ANOVA. By choosing a significance level of 5%, a three-stage sampling method was used to determine the sample and subsample size large enough for statistical conclusions to be valid. Statistics For each T% MPE (tolerance time) outcome measure, separate multivariate analysis of variance (MANOVA) was performed at each of the 18 time points. Contrast within the design matrix was used to determine if the affected lip side exhibits recovery for any of the groups and the relative rate of recovery among the

Fig. 11. Percent maximum possible effect is a mean to describe the recovery of thermal pain sensitivity. The x-axis shows the degree to which sensation is affected. One hundred percent indicates complete lack of sensation, 0% indicates complete return. Graph No. 1 shows complete sensory recovery in the sham group. Graph No. 2 shows minimal return after resection without repair in control animals. Graph No. 3, autogenous nerve graft, demonstrates approximately 85% return of thermal pain sensitivity. Graph No. 4, vicryl mesh, shows close to 70% recovery. Graph No. 5, vicryl mesh plus Matrigel demonstrates 80% recovery.

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Fig. 12. Morphometric examination of histologic sections allows determination of numerical and volume density of nerve fibers. The only differences between groups occurred in the middle sections. Volume density (dependent on size and number of fibers) is significantly decreased in the control, vicryl mesh, and vicryl mesh plus Matrigel groups when compared to sham. The autogenous group is not significantly different from sham. Numerical density (dependent on number of fibers only) parallels volume density. When examined independent of sham, there are no significant differences between the different graft methods. The asterisk indicates significant difference between groups at the P > .05 level.

groups. For histologic and morphologic data, a twoway ANOVA with the area, perimeter, number, and volume of the sample as the weighting factor was used to determine statistical significance based on surgical treatment group and time after surgery. When ANOVA indicated a statistical difference, Neuman-Keul’s multiple range tests were performed.

Results Thermal percent maximum possible effect Upon initial exposure to the thermal stimulus, responses varied tremendously. With repeated exposure, the animals eventually exhibited predictable and

Fig. 13. Surface density is determined by the use of a formula that examines perimeter (axonal and myelin) as a function of area. When axonal perimeter is used, only control and vicryl mesh groups are significantly different from sham. When myelin perimeter is used, all groups differ from each other. There are no differences among the different graft methods. The asterisk indicates significant difference at the P >.05 level.

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Fig. 14. Comparison of myelin perimeter, axon perimeter, and g-ratio (ratio of myelin size to axon size) shows considerable variability among groups. No significant differences between groups can be demonstrated.

reproducible responses. The presurgical sensory data showed a general reduction in the temperature at which the rabbits responded to the thermal stimulus over the course of five testing sessions with stable responses by the fourth sessions (Fig. 10). Postsurgical thermal stimulus responses varied depending on the treatment modality. The sham surgery group (Group 1) demonstrated no response to thermal stimulus immediately after surgery and began demonstrating preliminary stimulus response by postsurgical week 2 with 100% normal responses by 6 to 6.5 weeks (Fig. 11). The control/section without repair group (Group 2) demonstrated a complete anesthesia to stimulus for 5.5

to 6 weeks and then showed a small degree of recovery through the end of the study period. This group had approximately 10% recovery by week 9. The autogenous group (Group 3) demonstrated complete anesthesia initially and preliminary response to painful thermal stimuli at week 3 to 3.5, with approximately 85% recovery by week 9 (Fig. 12). The vicryl mesh group (Group 4) demonstrated preliminary thermal pain responses at week 5 and 60% to 70% return by the end of the study (Fig. 12). The vicryl mesh + Matrigel group (Group 5) demonstrated preliminary thermal pain responses at week 5 and 80% return of stimulus response by week 9.

Fig. 15. Photomicrographs comparing proximal, middle, and distal sections of sham and control groups. The darkly stained circular structures are myelin. Note the decreased number of myelinated nerve fibers in the middle section of the control (resection without repair) group (C) as compared to the sham group (S). Also note the pale staining of the nerve fibers distal to the graft.

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Fig. 16. Photomicrograph of representative sections taken from an IAN that is grafted using an autogenous nerve graft (A). Note the similarity to sham (S). The control/section without repair is shown (C). The myelinated nerve fibers in the autogenous group is smaller than in the sham group.

Significant differences were found between the thermal pain stimulus responses that existed before session ten and responses that were present after session ten for all groups except the control/section without repair. There was no statistical significant difference between the recovery of thermal pain stimulus response in the sham, autogenous, and vicryl mesh + Matrigel group.

Morphometric data In all analyses that involved density of nerve fibers, significant variability was noted only in the middle region (Fig. 12). Volume density, numerical density, and myelin surface density decreased significantly in the control/section without repair, vicryl mesh, and vicryl mesh + Matrigel groups when

Fig. 17. Photomicrograph of representative sections of an IAN in the vicryl mesh graft group (V), compared with the sham (S) and control/section without repair (C) groups. Note the close similarity of the middle section of the vicryl mesh to the control section. The distal section of the vicryl group contains dilated, pale staining, myelinated fibers.

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Fig. 18. Photomicrograph representative sections of an IAN grafted with vicryl mesh plus Matrigel (V + M) compared with sham (S) and control/section without repair (C) groups. A fascicular pattern present in the middle section of the V + M groups shows regenerating nerve fibers. The fibers are smaller than sham fibers in both proximal and middle sections.

compared to sham surgery group (Figs. 13 and 14). Axonal perimeter and area exhibited a large degree of variability in all groups except in the sham surgery and vicryl mesh-Matrigel groups, in which these

variables remained relatively constant from proximal to distal (Fig. 15). Axonal size decreased significantly in middle sections from the autogenous graft and the control/section without repair groups and increased

Fig. 19. Photomicrograph of mouse trigeminal ganglion demonstrating galactosidase staining in neuron cell bodies. An adenoassociated virus construct containing the lac Z gene is injected into the whisker pad of the mouse face. The vector is taken up by terminal endings of trigeminal neurons and retrogradely transported from the periphery to the more centrally located cell bodies, where the exogenous gene then integrated into the mouse DNA within the nucleus of the cell. The gene is subsequently expressed and galactosidase protein produced [67].

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relative to the sham group in distal sections taken from the control/section without repair group ( P > 0.01). Nerve fiber density decreased from all groups except sham and vicryl mesh + Matrigel ( P > 0.001). Axon numerical density ( P > 0.05) and fiber numerical density ( P > 0.01) were greater in the sham group than all the other groups. (Figs. 16 – 18) illustrate the light microscopic findings among the various surgical groups. Significant differences in the morphology of unmyelinated fiber populations occurred principally in the middle section rather than proximal or distal to the site of exposure, injury, or repair. Significant differences between the control/section without repair groups and all the other groups were found for surface density ( P > 0.005), surface area ( P > 0.05), and fiber perimeter ( P > 0.001). The vicryl mesh + Matrigel and vicryl mesh groups were significantly different from sham surgery for perimeter only ( P > 0.05).

Discussion The findings in this study concur with the hypothesis that autogenous nerve grafts, vicryl mesh, and vicryl mesh + Matrigel support IAN regeneration across a 1.1-cm defect in the rabbit. Regeneration across nerve gaps up to 1 cm in length in the rat model has been demonstrated when silicone tubes with a 1.2-mm diameter are used [63]. This model was chosen to represent a defect that would not heal spontaneously. The histologic and sensory data suggest that little nerve regeneration was evident in the control/section without repair group. The thermal sensory stimulus and response curves were based on an unanesthetized animal’s behavior and correlate to nociceptive somatosensation in animals and humans. The results indicated that vicryl mesh + Matrigel combination supports the regeneration of sensory axons that provide recovery of nociceptive information. The regeneration and recovery of nociceptive somatosensation was equal to that obtained by autogenous grafting in the same model. Vicryl mesh without Matrigel supported nerve growth but did not provide the same degree of regeneration to provide nociceptive somatosensation compared to the autogenous graft or vicryl mesh + Matrigel graft. The histologic and morphometric analyses demonstrated that autogenous, vicryl mesh, and vicryl mesh + Matrigel grafts support sensory nerve regeneration across a nonhealing IAN defect in the rabbit. Although not statistically significant, the autogenous graft group tended to demonstrate superior fiber regeneration than the other groups. The combination

of vicryl mesh + Matrigel provided more nerve regeneration than did vicryl mesh without Matrigel. It seems that vicryl mesh with intraluminal Matrigel and, to a lesser extent, vicryl mesh alone supports nerve regeneration. When compared with autogenous grafts, the biodegradable alloplastic nerve grafts performed well. The neurotropic properties of the solubilized basement membrane, Matrigel, allow the regenerating nerve fibers to gap the defect quicker and to a greater degree than in its absence. It is hypothesized that the physical and neurotropic properties of Matrigel preserve luminal patency and enhance or guide nerve sprouts during regeneration while reducing the scar tissue infiltration through the vicryl polyglactin mesh. Basement membranes form a supporting structure for many cells in the body, including muscle fibers and Schwann cells. The predominant basement membrane components that are found in Matrigel include type IV collagen, laminin, and the proteoglycans. Different isoforms confer cell- or tissue-specific functional and behavioral activity [46]. Specialization of these proteins in different cells provides an appropriate structural framework or scaffold on which neurons grow, for instance. With regard to the nervous system, the basement membrane or basal lamina produced by Schwann cells lines the endoneurial tubes and provides a supportive environment in which elongating, regenerating axons can grow [21,61]. The laminin in the basal lamina of the endoneurial tubes is similar to the basement membrane laminin of muscle fibers [66]. Specially treated muscle grafts can support regeneration just as well as peripheral nerve grafts [55,56]. The results demonstrated in the current study show that even in the absence of an intact structural framework (i.e., no endoneurial tubes or muscle fiber scaffold), the extracellular matrix laminin proteins within Matrigel support peripheral regeneration. Other than the basement membrane proteins mentioned previously, Matrigel does not have neurotropic factors or contain living cells that produce the more conventional neurotrophic factors such as nerve growth factor [27]. Inferior alveolar nerve regeneration was similar in the vicryl mesh + Matrigel group and autogenous graft groups, however. In the autogenous graft groups, the Schwann cells remained viable and it seems reasonable to assume that these Schwann cells provided for enhanced regeneration [2,3,49]. It is likely that in the vicryl mesh + Matrigel group, however, the Schwann cells migrated from the proximal nerve stump and guided the regenerating axon [58] to the distal stump that maintains Schwann cell – lined endoneurial tubes.

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The importance of Schwann cells has been confirmed in an experimental setting in which Schwann cells are eliminated and subsequent peripheral axon regeneration is reduced [13]. Schwann cells produce neurotrophic factors that may enhance the ability of a nerve to regenerate [3,22,59]. Not only do Schwann cells produce trophic factors, but so do target tissues that the axons attempt to innervate after injury [19,25]. For instance, the skin may chemotropically guide axons toward it and then trophically maintain those axons that reach it [30]. The production of many factors that have been shown to enhance regeneration are upregulated after peripheral nerve damage, including factors such as nerve growth factor itself, neurotrophin 3, neurotrophin 4, neurotrophin 5, brain-derived neurotrophic factor, ciliary neurotrophic factor, platelet-derived growth factor, fibroblast growth factor, glial growth factor, and glial-derived growth factor [51,57]. These factors can concentrate at the site of nerve injury within a synthetic chamber [33]. Based on the evidence demonstrated in this study, even within the confines of a biodegradable sheath, endogenous factors (and basement membrane proteins) may suffice to yield an anatomic and functional nerve. This may have been accomplished if neurotrophic factors permeated into the membrane or Matrigel or, as in the case of Schwann cells, progressed with the growing axon. It is likely that some combination of neurotrophic factors and the basement membrane proteins in Matrigel acted in conjunction to give the outcome observed in this study. Previous research has demonstrated that the addition of exogenous neurotrophic growth factors enhances peripheral nerve regeneration [20,52]. It is probable, however, that a more complex manipulation of cell types and growth factors is required to enhance regeneration optimally in instances in which the

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nerves are severely injured. Within the trigeminal nervous system, subsets of sensory neurons depend on different growth factors and receptors for their development and survival. Nerve growth factor and its receptor are associated with A-delta and C fibers. Neurotrophin 4 and brain-derived neurotrophic factor support the larger diameter A-beta fibers [1]. Aside from these, the expression of neurotrophin 3, neurotrophin 5, and ciliary neurotrophic factor have been localized within the trigeminal system and target tissues, although their functional specificity remains unclear [7 – 9,24,48]. The clinical implication of this study was that biodegradable alloplastic nerve grafts potentially could be used to repair trigeminal nerve injuries in humans. Based on animal data, this technique should yield results comparable to those obtained with autogenous human nerve grafts. Like previous clinical studies, entubulation supports nerve regeneration and neurotropic substances enhance it. This study supports the hypothesis that the combination is synergistic and the future application of entubulation and neurotropic substances to injured human nerves should be beneficial. Future research should involve the use of other allogeneic graft materials with easier handling properties. A major shortcoming of the vicryl mesh method was that a sheet being formed into a tubular prosthesis was a tedious technique. The extra manipulation of the nerve for tubulation of the material may cause additional ingrowth of fibrous connective tissue and add to the amount of trauma. The preformed corrugated polyglycolic acid tubes described by MacKinnon and Dellon [34] may be a good alternative. We must continue our search for the ideal nerve graft. If an allogeneic material is useful, it must (1) permit rapid reestablishment of blood supply, (2) be

Fig. 20. Photomicrograph of lac Z gene transfer into mouse facial skeletal muscle. Both subcutaneous muscle and muscle cells around hair follicles are transduced to express a foreign (nonmouse) reporter gene. Masticatory muscles (i.e., masseter) can also be genetically manipulated (data not shown).

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nonreactive and cause minimal scar tissue formation, (3) provide a containment mechanism for guided regeneration of nerve fibers, (4) confine neurotropic and neurotrophic factors to allow axonal chemotaxis, and (5) be technically and economically feasible.

Oral and Maxillofacial Surgery at the University of North Carolina at Chapel Hill. The publication of this article should contribute to the known literature and encourage future research into an important clinical problem. We are honored to have been collaborators of Dr. Hegtvedt in his efforts to enhance our clinical and basic science understanding of our profession.

Future directions The use of biodegradable materials and the local application of exogenous neurotrophic factors may prove valuable in the future for the treatment of trigeminal nerve injury. Currently, the use of virally mediated gene therapy strategies for treating various neuromuscular disorders is being contemplated as an alternative means to enhance peripheral nerve regeneration. Original research by one of the co-authors has demonstrated that an exogenous gene can be inserted into the mouse genome for long-term expression of a reporter gene (lac Z). In neurons and muscle cells, the lac Z gene (derived from Escherichia coli DNA), which encodes for the beta galactosidase protein, is expressed (Figs. 19 and 20) [50]. The use of a reporter gene allows an investigator to determine if gene transfer occurs and in what tissues the gene is expressed and protein produced. The logical extrapolation of this is to insert into a viral vector neurotrophic genes or other genes that either enhance neuron survival or promote axonal outgrowth after injury. The manipulation of neurons might enhance nerve regeneration. Muscle tissues (or epithelial tissue) within the distribution of an injured nerve might be manipulated to serve as local ‘‘factories’’ of regeneration promoting substances. Direct injections or local application of a viral vector is one means by which to attempt this methodology; however, carrier materials (e.g., Matrigel within a biodegradable sheath) also might be loaded with vector (or just the trophic factors alone) to maintain a more precise and prolonged delivery to nearby tissues. Vectors can be designed to be tissue-specific in regard to their infectivity and their temporal production of a desired factor. Such biotechnology is likely to become valuable in the near future for more than just the treatment of peripheral nerve injuries.

Acknowledgment This article is dedicated to Dr. Arden Hegtvedt in memory of his accomplishments and contributions to our profession. This study was performed by Dr. Hegtvedt in partial fulfillment of the requirements for the degree of Masters of Science in the Department of

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Index Note: Page numbers of article titles are in boldface type.

A A-2186, in prosthodontic rehabilitation, 78 Acrylic resins, in prosthodontic rehabilitation. See Prosthodontic rehabilitation. Adhesive tapes, in wound closure, 100 – 101 Alginate carriers, uses of, 35 – 36 AlloDerm as dermal substitute, 65 as oral mucosa substitute, 69 – 71 cosmetic uses of, 57 Allografts as dermal substitutes, 65 cosmetic uses of, 56

for craniofacial defects, 9 historical aspects of, 3 – 4 human studies of, 10, 12 signaling mechanisms in, 4 – 8 activity regulation in, 6 family characteristics in, 4 – 6 gene expression in, 7 – 8 serine/threonine kinase activity in, 6 Smads in, 6 – 7 Bone regeneration, guided. See Guided bone regeneration. Bovine pericardium, in guided bone regeneration, 26

Alloplast, cosmetic uses of, 55 Apligraf, cosmetic uses of, 56 Artecoll, cosmetic uses of, 55

C Calcium sulfate barrier, in guided bone regeneration, 26

Autologous chondrocytes, in cartilage regeneration, 110-112

Calthane, in prosthodontic rehabilitation, 79 – 80

Autologous fibrin tissue adhesive, in wound closure, 102

Cartilage, cosmetic uses of, 60

B Biodegradable materials, for peripheral nerve injuries. See Peripheral nerve injuries.

Cartilage regeneration, 107 – 118 articular cartilage architecture and, 108 calcification in, 108 – 109 problems with, 107 tissue engineering in, 108 – 115 allografts and ex vivo cartilage, 110 – 112 for full-thickness defects, 109 – 110 for osteochondral defects, 113 – 115 for partial-thickness defects, 109 need for cells in, 109

Bio-Guide, in guided bone regeneration, 25 BioMend Extend, in guided bone regeneration, 25 – 26 Bioresorbable barriers, in guided bone regeneration. See Guided bone regeneration. Bone, cosmetic uses of, 57 – 59 Bone morphogenetic proteins, 3 – 16 after bone injury, 8 – 9 animal studies of, 9 – 10

Capset, in guided bone regeneration, 26

Chemical polymerization, of in situ forming biomaterials, 37 Cleft palate, prosthodontic rehabilitation for. See Prosthodontic rehabilitation. Collagen, in guided bone regeneration, 25 – 26 Cosmesil, in prosthodontic rehabilitation, 78

1042-3699/02/$ – see front matter D 2002, Elsevier Science (USA). All rights reserved. PII: S 1 0 4 2 - 3 6 9 9 ( 0 2 ) 0 0 0 3 9 - 0

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Index / Oral Maxillofacial Surg Clin N Am 14 (2002) 133–136

Cosmetic materials, advances in, 55 – 61 bone, 57 – 59 cartilage, 60 soft tissue, 55 – 57 Craniofacial defects, bone morphogenetic proteins for, 9 Cyanoacrylates, in wound closure, 101 – 102

D Decalcified freeze-dried bone allografts, in guided bone regeneration, 27

future goals of, 27 – 28 historical aspects of, 17 – 18 membrane design criteria in, 21 – 22 nonresorbable barriers in, 22 – 24 osteopromotion principle in, 19 – 20

H Hard tissue replacement, in craniofacial reconstruction, 58

Dermalogen, cosmetic uses of, 56

I Implant-retained maxillofacial prostheses, studies of, 88 – 90

Dressings, in wound closure, 102 – 103

Implants, endosseous. See Endosseous implants.

Dura mater, in guided bone regeneration, 26

In situ forming biomaterials, 31 – 40 chemical polymerization of, 37 diffusion of, 34 – 35 enzymatic cross-linking of, 36 – 37 ionic cross-linking of, 35 – 36 pH triggered systems, 34 photopolymerization of, 37 – 38 self-assembling systems, 38 swelling of, 34 thermally triggered systems, 31 – 34

Dermagraft, as dermal substitute, 65

E Endosseous implants, 41 – 53 biologic modifications of, 47 – 48 macroretentive features of, 42 microretentive features of, 42 – 47 alterations of surface oxide, 46 – 47 biologic response, 45 – 46 surface roughness, 42 – 44 by blasting or etching, 44 – 45

Integra, as dermal substitute, 65

Engineered tissue, cosmetic uses of, 56 – 57 Epicel, as dermal substitute, 64

L Lambone, in guided bone regeneration, 26

Epithane-3, in prosthodontic rehabilitation, 79 – 80

Laminar bone, in guided bone regeneration, 26 Leucine zipper motif, characteristics of, 38

F Fibrin glue, in wound closure, 102

G Gore-Tex cosmetic uses of, 55 in guided bone regeneration, 23 – 24 Growth factors, in bone healing, 59 Guided bone regeneration, 17 – 29 biology of, 20 – 21 bioresorbable barriers in, 24 – 27 natural products, 25 – 26 synthetic products, 26 – 27 clinical application of, 27 experimental studies of, 18 – 19

Liposomes, uses of, 33 – 34

M Medpore, cosmetic uses of, 57 Mesenchymal stem cells, in cartilage regeneration, 112 Millipore membrane filters, in guided bone regeneration, 24 Myverol, uses of, 32

N Nonresorbable barriers, in guided bone regeneration, 22 – 24

Subject Index / Oral Maxillofacial Surg Clin N Am 14 (2002) 133–136

O Obturator prostheses, in prosthodontic rehabilitation. See Prosthodontic rehabilitation. 2-Octylcyanoacrylate, in wound closure, 101 – 102

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Polylactic acid, in guided bone regeneration, 26 Polymethacrylic acid, uses of, 34 Polytetrafluoroethylene, in guided bone regeneration, 23 – 24, 27

Oral mucosa substitutes, 67 – 71 in vitro culture of, 67 – 68 tissue-engineered, 68 – 71 versus skin, structure and function of, 67

Polyurethanes, in prosthodontic rehabilitation, 79 – 80

Osseoquest, in guided bone regeneration, 27

Prosthodontic rehabilitation, 75 – 95 acrylic resins in, 80 – 84 coloration of, 80 – 81 extrinsic, 82 intrinsic, 81 – 82 copolymer, 80 emulsions of, 83 – 84 light-cured, 80 methyl methacrylate, 80 processing of, 80 retention of, 82 – 83 adhesives in, 83 pressure-sensitive tape in, 83 silicone adhesives in, 83 constraints in, 75 – 76 economic aspects of, 90 – 91 for cleft palate, 84 – 86 active appliances in, 85 orthopedic appliances in, 84 – 85 passive appliances in, 85 pin-retained appliances in, 85 presurgical nasoalveloar molding in, 85 – 86 future directions in, 91 implant-retained maxillofacial prostheses in, 88 – 90 extraoral, 89 – 90 intraoral, 88 – 89 obturator prostheses in, 86 – 88 definitive, 86 – 87 interim, 86 studies of, 87 – 88 surgical, 86 palatal augmentation prostheses in, 88 palatal lift prostheses in, 88 polyurethanes in, 79 – 80 silicones in, 76 – 79 foam, 78 HTV silicone, 78 – 79 polymers, 78 room temperature vulcanizing silicones, 77 studies of, 77 – 78 wettability of, 78 speech aid prostheses in, 88 versus surgery, 76

Osteochondral defects, cartilage regeneration in, 113 – 115 Osteopromotion principle, in guided bone regeneration, 19 – 20

P Palamed, in prosthodontic rehabilitation, 80 Palatal augmentation prostheses, indications for, 88 Palatal lift prostheses, indications for, 88 Peripheral nerve injuries, biodegradable materials for, 119 – 134 animal studies of, 120 – 121 future directions in, 131 – 132 postsurgical sensory training in, 123 – 124 presurgical sensory training in, 121 – 122 results of, 124, 126 – 130 morphometric data, 127 – 130 thermal percent maximum possible effect, 124, 126 – 127 statistics on, 124 surgical procedures in, 122 – 123 Photopolymerization, of in situ forming biomaterials, 37 – 38 Pluronics, uses of, 32 Poly-l-lactic-co-glycolic acid, uses of, 32, 59 Poly-N-isopropyl acrylamide, uses of, 32 Polyacrylic acid, uses of, 34 Polyethylene glycol photopolymerization of, 37 uses of, 34 Polyethylene oxide, uses of, 32 Polyglactin 910 mesh, in guided bone regeneration, 26 Polyglycolic acid, in guided bone regeneration, 26

Porcine collagen membrane, in guided bone regeneration, 25

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Index / Oral Maxillofacial Surg Clin N Am 14 (2002) 133–136

R Resolute XT, in guided bone regeneration, 26 – 27

T Tissue adhesives, in wound closure, 101 – 102 Trismus, and fabrication, of obturator prostheses, 87

S b-Sheet structures, characteristics of, 38 Silastic, in prosthodontic rehabilitation, 78

V Vicryl, in guided bone regeneration, 26

Silicone prostheses, in prosthodontic rehabilitation. See Prosthodontic rehabilitation. Siphenylenes, in prosthodontic rehabilitation, 78 Skin grafts, cosmetic uses of, 56 Skin substitutes, 63 – 67 bilayers of epithelium and dermis, 65 – 67 dermis, 64 – 65 epithelium, 64 Soft tissue, cosmetic uses of, 55 – 57 Speech aid prostheses, indications for, 88 Speech intelligibility, obturator prostheses and, 87 Staples, in wound closure, 100 Sutures, in wound closure, 99 – 100

W Wound closure materials, 97 – 106 adhesive tapes, 100 – 101 dressings, 102 – 103 future directions in, 103 – 104 historical aspects of, 97 – 98 patient assessment for, 78 postoperative care and, 103 staples, 100 sutures, 99 – 100 tissue adhesives, 101 – 102 wound anesthesia and, 78 wound assessment for, 78 wound preparation for, 98 – 99