Ординатура / Офтальмология / Английские материалы / Biomaterials and regenerative medicine in ophthalmology_Chirila_2010

How should a pegged hydroxyapatite implant be classified: buried integrated or exposed integrated? Perry (1991) suggested that eventual vascularization of the drilled hole would allow re-epithelialization, permitting the motility advantage of exposed integrated implants with the safety profile of a buried implant. In practice, however, drilling may become complicated by exposures, chronic discharge and formation of pyogenic granulomas at the peg site. In one series the complication rate of pegging was 38% (Jordan et al. 1999a). The most common problems in this review (n = 165) included: chronic discharge, 37%; pyogenic granuloma formation, 31%; peg extrusion, 29%; poor transfer of movement, 11%; clicking with extreme gaze, 11%

(the clicking appears to be related to a loose fit between the peg and ocular prosthesis, such that, when the ocular prosthesis is restricted in its motion by the fornices in extreme gaze, the peg continues to travel with the implant and knocks against the side of the drilled shaft in the implant; drilling a shaft with smaller diameter may prevent this problem). Peg complication rates as high as 67% (n = 275) have been reported (Shoamanesh et al. 2007). A study of complications associated with freestanding polycarbonate pegs found a complication rate of 71% (n = 21) (Fahim et al. 2007).

There are a number of available peg systems. In early polycarbonate pegs, used with the hydroxyapatite implant, the peg was attached to the ocular prosthesis, with the female end drilled into the implants. This made insertion and removal of the ocular prosthesis difficult and potentially traumatic to the conjunctival lining (Edelstein et al. 1997; Oestreicher et al. 1997). Thus a permanent peg was placed inside the drilled hole, with a ball at the exposed end to articulate with a corresponding indentation carved into the posterior surface of the ocular prosthesis (Perry 1991; Shields et al. 1994).

To decrease the high complication rate associated with non-sleeved PMMA and polycarbonate pegs, titanium peg and sleeve systems have been advocated (Jordan & Klapper 2000). Studies suggest that peg extrusion rates and pyogenic granuloma formation are reduced with sleeved peg systems (Edelstein et al. 1997; Jordan & Klapper 2000; Lee et al. 2002). In one series, sleeved peg extrusion rates were 11% (n = 74) compared with 27% (n = 191) for non-sleeved pegs (Lee et al. 2002). A recent large retrospective review (n = 353) found a significantly lower incidence of peg extrusion and granuloma formation with titanium pegged implants as compared with non-sleeved PMMA and sleeved polycarbonate pegs (Yoon et al. 2008).

Primary placement (i.e. at the time of enucleation) of a titanium sleeve in porous implants has been suggested to reduce the cost and complication rates (in theory) of secondary drilling and peg placement (Liao et al. 2005a; Liao et al. 2005b). In a study of 52 patients receiving a primary peg placement, a 29% complication rate was found. This complication rate is similar to reported outcomes of secondary peg insertion (Yazici et al. 2007). Recent

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surveys suggest there is a decline in placement of motility pegs by surgeons (Su & Yen 2004; Viswanathan et al. 2007).

17.5.2 Porous polyethylene implants

Despite the initial success of hydroxyapatite and reports of low extrusion rates when wrapped in donor sclera, a number of problems persist:

(a)a theoretical infectious risk associated with donor sclera (e.g. prion disease – although we are not aware of any reports to date);

(b)development of late exposures and pyogenic granuloma formation (Fig. 17.8);

(c)difficulty in re-drilling in cases of implant migration.

Some of these problems have been addressed by a new generation of porous polyethylene (Medpor®) implants, which can be placed safely without wrapping. An exposure rate of 1% (3/302) was reported for Medpor® implants placed without wrapping (Chen & Cui 2006). Whereas the coralline hydroxyapatite implant must be drilled with a power-tool for peg placement, porous polyethylene may be pegged by hand with a screw driver. When a hydroxyapatite sphere rotates in the orbit, re-drilling may create a large tract that does not support a peg well. With porous polyethylene, repositioning the peg is less problematic, as the tract left by the prior screw is a narrow spiral.

Like hydroxyapatite, porous polyethylene permits fibrovascular ingrowth

(Karesh & Dresner 1994; Rubin et al. 1994). Porous polyethylene became available for orbital implantation in 1991 (Porex Surgical Inc., Newnan, GA). Animal model studies suggest that porous polyethylene incites less inflammation and fibrosis than hydroxyapatite (Goldberg et al. 1994; Li et al. 2001). Using electron microscopy, porous polyethylene implants show a smoother surface than hydroxyapatite (Bio-Eye®), synthetic hydroxyapatite (FCI3®) and aluminum oxide (alumina) implants (Jordan et al. 2004). The rate of vascularization of porous polyethylene appears to be slower than that for hydroxyapatite (Bio-Eye®), synthetic hydroxyapatite (FCI3®), and aluminum oxide implants (Jordan et al. 2004). Porous polyethylene implants with 400 um pore size vascularize more rapidly than the 200 um pore size (Goldberg et al. 1994; Rubin et al. 1994). Implant vascularization appears to be faster with the Medpor-Plus® implant, which is a combination of Medpor® and synthetic bone graft particulate (Novabone®) in a 70:30 ratio (Naik et al. 2007).

The first generation of spherical porous polyethylene implants had a rough surface (Li, et al. 2001). Subsequently, implants with a smoother anterior surface were introduced (Woog et al. 2004). A retrospective report of 91 enucleations suggested similar exposure rates for wrapped and unwrapped

porous polyethylene implants (<5%) (Blaydon et al. 2003).

Medpor® implants currently come in a number of shapes. These include but are not limited to:

(a)simple sphere;

(b)conical implants with a flat 6 mm anterior surface (Dresner);

(c)smooth surface tunnel (SST™ sphere) with suture tunnels for easier muscle attachment (Woog et al. 2004);

(d)conical orbital implant (Rubin et al. 1998) which incorporates a superior projection to reduce superior sulcus defect;

(e)‘Quad’ motility implant (Anderson) – similar in design philosophy, shape and method of muscle attachment (imbrication) to the Iowa and Universal implants (Anderson et al. 1990; Anderson et al. 2002).

17.5.3 Porous aluminum oxide (Al2O3) implants

Implants using aluminum oxide (also known as alumina) are a recent instalment in the continuing search for better-tolerated orbital implants. Aluminum oxide has been in use for over 30 years as an implant in orthopedics and dentistry (Smith 1963). Osteoblasts and fibroblasts appear to grow faster on aluminum oxide than on hydroxyapatite in laboratory models (Mawn et al. 2001). Reports of aluminum encephalopathy with the use of ionocem (a biomaterial made by reacting calcium aluminum fluorosilicate with polyalkenoic acid), have raised concerns regarding aluminum-containing biomaterials. However, aluminum oxide appears to be bioinert and insoluble in tissues. Blood samples drawn from patients with alumina implants show normal aluminum levels (Christel 1992; Jordan et al. 2000c).

Outcomes in a review of 107 alumina implants placed over 3 years (all wrapped in polyglactin mesh) were encouraging (Jordan et al. 2003a). Of these, 107 implants, 76 were enucleation implants (50 as secondary implants). The overall exposure rate was 2% in 107 implants (0 in the enucleation group), as compared with 3% in 120 FCI3® implants (1995–1999), and 8% in 258 Bio-Eye® implants (1990–1995) for the same authors. However, a recent retrospective review of 108 patients found a long-term exposure rate of 7.4% for alumina implants (mean follow-up of 35.8 months) (Wang et al. 2009). We will have to await long-term data to make firm conclusions about the clinical efficacy of aluminum oxide as an orbital implant.

17.6Trends in pediatric enucleation

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tumor recurrence (Shields et al. 1992b; Christmas et al. 2000) – since a significant portion of pediatric enucleations are done for retinoblastoma; and (b) less disfiguring appearance of acquired anophthalmia (with ocular prosthesis only) in young children as compared with adults – at least in the short term. However, animal model studies support the theory that placement of an implant is necessary for stimulation of orbital growth (Cepela et al. 1992). Studies on the long-term effects of pediatric enucleation in the four decades spanning 1935–1975, showed stunted orbital growth in children who did not receive implants (Taylor 1939; Pfieffer 1945; Apt & Isenberg

1973; Osborne et al. 1974). These observations have prompted the routine use of orbital implants after pediatric enucleation. Human follow-up studies of retinoblastoma cases suggest that childhood enucleation, when combined with a large implant, minimizes orbital growth retardation (Fountain et al. 1999). In addition, the role of the ocular prosthesis in stimulating orbital growth cannot be ignored (Yago & Furuta 2001).

In normal development, by 5 years of age orbital volume has reached about 80% of the volume seen at 15 years, in both sexes (Bentley et al. 2002). Earlier reports suggested attainment of 80% of adult size by 3 years of age (Scott 1954). Orbital volume is thought to reach adult volume by 12–14 years of age (Yago & Furuta 2001). When enucleation takes place in infancy, placement of progressively larger implants has been advocated to achieve adequate orbital growth (Vistnes 1987) In animal models, implantation of a sphere that inadequately compensates for volume loss, does not stimulate orbital growth (Sarnat & Shanedling 1970; Sarnat & Shanedling 1972; Sarnat & Shanedling 1974; Sarnat 1979; Sarnat 1981; Sarnat 1982; Reedy et al. 1999). In a recent study of orbital volume following childhood enucluation for retinoblastoma, 3 of 13 hydroxyapatite implants were noted to have migrated. In all 3, the orbital volume on the ipsilateral side was found to be larger than on the non-operated side. In the absence of implant migration, the operated side had a smaller volume in all cases.

This argues for the significance of mechanical stimulation in orbital growth

(Lyle et al. 2007).

Although stunted orbital growth following enucleation is well established in radiographic studies, there may not be obvious cosmetic facial asymmetry (Howard et al. 1965; Hintschich et al. 2001). More recently, orbital volume measurements using computed tomography have suggested that enucleation, in both children and adults, is associated with reduction of bony orbital volume over time (Hintschich et al. 2001). Other studies have found a greater impact of radiation dose (megavoltage external beam irradiation) on orbital growth than implant placement and size. In a study of irradiated orbits, secondary enucleation did not appear to have an additive growth-retarding effect (no implant placed) (Imhof et al. 1996). Orbital growth appears to be most affected when radiation is given in the first year (particularly in the first 6

months) of life (Imhof et al. 1996; Ameniya et al. 1977; Peylan-Ramu et al. 2001).

Do these studies imply that the placement of an implant after enucleation is unnecessary? No. The wealth of human clinical and experimental animal studies showing the importance of volume replacement in promoting orbital growth (Taylor 1939; Pfieffer 1945; Kennedy 1964; Apt & Isenberg 1973;

Osborne et al. 1974; Fountain et al.1997; Yago & Furuta 2001; Chen & Heher 2004) cannot be discounted. Post-enucleation socket syndrome is still an important consideration and an excellent argument for placement of an implant of adequate volume in any patient. In combination with a buried integrated implant, an ocular prosthesis can show nearly life-like movements to match those of the contralateral eye, an important aspect of rehabilitation.

There has been an increasing trend to use porous implants in children (De Potter et al. 1992; De Potter et al. 2004). Fibrovascular ingrowth into these implants makes later removal difficult. (Kaltreider et al. 2001). One way to achieve implant motility and at the same time permit later implant exchange is to use an acrylic ball and attach the muscles to the conjunctival fornices (Soll 1972; Nunery & Hetzler 1983). A potential problem with the use of hxdroxyapatite for such an implant is that hydroxyapatite implants are radio-opaque on imaging. In theory this could interfere with detection of calcification associated with tumor recurrence in cases of retinoblastoma.

However, the well-circumscribed appearance of the implant on imaging, and its intermediate signal intensity, are thought to be characteristic enough to not interfere with the radiologic features of retinoblastoma recurrence (De Potter et al. 1992; De Potter et al. 1994). In addition, coralline hydroxyapatite implants do not appear to attenuate external beam photon radiation significantly (Arora et al. 1992). When calcification is detected following enucleation with placement of a scleral-wrapped orbital implant, the presence of calcium may be dystrophic, and does not necessarily indicate recurrent tumor growth. (Summers 1993).

An autogenous dermis fat graft is another option (Bosniak et al. 1989). Graft atrophy, usually more pronounced when dermis fat grafts are placed in older patients, may also occur in children. Dermis fat graft hypertrophy may occur in growing children (Guberina et al. 1983; Mitchell et al. 2001)

In the first few months after enucleation, the rough surface of porous implants may be associated with an even higher exposure risk in children than in adults. Although there are only a handful of studies in children, scleral wrapped hydroxyapatite and polyethylene implants appear to have a reduced exposure rate compared with unwrapped implants (De Potter et al. 1994; Karcioglu et al. 1998; Christmas et al. 2000; Lee et al. 2000; Nolan et al. 2003; Iordanidou & De Potter 2004). The reasons for this are not well established.

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17.7Gaps in scientific knowledge and future trends

Despite the 50+ years of evolution of motility implants, our understanding about the actual motility of various implants, and the efficacy of pegging in improving motility is mostly subjective – prone to the bias of optimistic patients and surgeons. The major determinants of implant and ocular prosthesis motility are unclear. An important question is whether ocular prosthesis motility is mainly due to the retraction of conjunctival fornices by muscle contraction (De Voe 1945; Coston 1970; Soll 1982; Nunery & Hetzler 1983; Tyers & Collin 1985; Smit et al. 1991a), or is due to direct transmission of forces between implant and ocular prosthesis (Soll 1972; Nunery & Hetzler 1983)? In a case series of 25 patients, Nunnery and Hetzler showed that direct suturing of the rectus muscles to the conjunctival fornices can impart adequate motility to a ocular prosthesis (Nunery & Hetzler 1983).

Adequate ocular prosthesis motility does not imply movement to fully match the normal side. This would create the problems with edge show and torsion that were seen with the Allen implant (Jordan et al. 1987). Adequate implant motility is probably about 70–75% of the contralateral normal eye. Still, a major hurdle to social rehabilitation after enucleation continues to be poor apparent ocular prosthesis movement. Apparent, since it is unclear whether the problem is primarily an issue of inadequate implant movement, or poor translation of movement to the ocular prosthesis. Alternatively, the ocular prosthesis may be limited in motion by the conjunctival fornices, with the implant slipping underneath (Nerad et al. 1991). Could a large ocular prosthesis, within the confinesofasmallpalpebralfissureandconjunctivalfornices,restrictmovement of the implant – especially when a peg is placed?

Studies on the relationship between implant and ocular prosthesis motility are generally lacking. A handful of investigators have attempted to assess the efficacy of various motility implants objectively; however, there are significant shortcomings. One simple approach to measuring implant motility has been to mark the conjunctival center and measure excursions in millimeters (Bosniak et al. 1989; Custer et al. 1999; Long et al. 2003) Custer and associates compared implant motility of scleral-wrapped acrylic (n = 7), silicone (n = 8) and hydroxyapatite (n = 31) implants. They found similar implant motility. Only age and implant size were found to be significant predictors of implant movement: vertical and horizontal movement decreased with increasing age and increased with larger implant sizes (Custer et al. 1999). Bosniak and associates compared implant motility between synthetic spherical implants (n = 47, all with muscle imbrication, implant sizes 14–20 mm) and autogenous dermis fat grafts (n = 34).(Bosniak et al. 1989). Overall they found better implant motility of dermis fat grafts. The diameter of spherical implants was not found to be a significant predictor of implant movement. Using the temporal and nasal limbus of the ocular prosthesis as

a landmark for comparison, the difference in motility between dermis fat and synthetic implants disappeared. The authors found no correlation between ocular prosthesis motility and forniceal depth (Bosniak et al. 1989). Another, similar, approach is to use Kestenbaum spectacles to measure ocular prosthesis motility (Smit et al. 1991a). Using this approach, Smit and associates found no significant difference in ocular prosthesis motility between Allen (n =

12) and primary baseball (sclera-covered 18 mm acrylic, n = 15) implants. Motility of the ocular prosthesis was lowest in patients with no implant (n = 11). Patients with secondary baseball implants (n = 11) demonstrated ocular prosthesis motility in between the primary implantation group and patients with no implant.

Using a straight ruler to measure displacement on the curved surface of implants and ocular prostheses is inaccurate. Such measurements are subject to a significant cosine function error and are not comparable for spheres of different size, unless millimeters of movement are translated into degrees of rotation. Steven’s hypothesis does not hold true for prosthetic eyes (Steven’s hypothesis assumes that, for an emmetropic eye, the center of rotation and each end of the corneal diameter (12 mm) form an equilateral triangle; on this basis, linear displacement of the ocular limbus, as measured by a Wessely keratometer, can be converted into degrees of rotation) (Yamashiro 1957).

A more sophisticated approach for measuring ocular prosthesis movement is the magnetic search coil technique (Collewijn et al. 1975): the current generated by movement of a wire coil (placed on the ocular prosthesis) in a uniform magnetic field is used to derive the amplitude and velocity of movement (Nerad et al. 1991; Colen et al. 2000). In a series of 16 patients with Iowa and Universal implants, Nerad and associates found good initial transmission of movement to the ocular prosthesis, but final amplitudes that were on average about 50% of target movement (target = 10 degrees rotation) (Nerad et al. 1991). Colen and associates found similar prosthetic eye saccadic amplitudes between scleral wrapped spherical acrylic (n = 16) and hydroxyapatite (n = 14) implants (Colen et al. 2000). Unfortunately, the magnetic search coil technique does not give information about implant motility. It would be interesting to use this method to evaluate movement in the same patient, before and after placement of a peg. More recently, infrared oculography has been used in a prospective case series of 10 patients to quantitatively assess the effect on ocular prosthesis motility of placement of a peg. Although vertical motility was not significantly affected by peg placement, horizontal excursions did improve (Guillinta et al. 2003). Lucci and associates used a prism in front of the normal eye and a fixed target at 1 meter to measure ocular prosthesis movement. The prism power was increased until there was no more improvement in ocular prosthesis motility. Using this approach, a similar amplitude of movement was found in spherical and quad (Medpor®) implants (Lucci et al. 2007).

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The development of an objective system to evaluate implant and ocular prosthesis motility is important for further development and refinement of motility implants. Such a system should permit simultaneous comparison of implant and ocular prosthesis movement, allowing investigators to better understand the dynamic relationship between the implant and ocular prosthesis movement. The surgical techniques and materials used in implantation may be culprits in limiting adequate implant motility. For example: (a) does imbrication of the rectus muscles over the surface of spherical implants produce a motility restriction by over-stretching the medial and inferior recti? (Dr A. Jampolsky, San Francisco, CA, personal communication); (b) Could the use of porous implants actually limit implant motility in the long run, in that fibrovascular ingrowth of the muscle bellies into the implant creates an adhesion-type syndrome in some patients?

To better illustrate the importance of observing ocular prosthesis and implant motility simultaneously, a simple experiment may be performed. The original impression mold of the socket may be used to produce a clear acrylic ocular prosthesis of the same shape, surface irregularities and frictional properties as the patient’s own white acrylic ocular prosthesis. This allows both the socket and the implant to be observed during movement. Placement of a central conjunctival mark (with a tissue marker) allows the observer to better appreciate the relative position of implant and ocular prosthesis. This is illustrated in Fig. 17. Note that, in Fig. 17.10, poor ocular prosthesis abduction is associated with poor implant abduction, but the poor ocular prosthesis adduction is not associated with poor implant movement. The implant has slipped underneath the ocular prosthesis, suggesting that the limitation of ocular prosthesis adduction is probably related to forniceal restriction to movement. Both the implant and ocular prosthesis appear to elevate well. In depression, the ocular prosthesis appears to move well, masking the inferonasal rotation of the implant that is appreciated through the clear ocular prosthesis.

The search for more biocompatible orbital implants continues. In particular, the relationship between the microsturcture of porous implants and clinical performance requires further investigation. The ideal orbital enucleation implant would have a similar weight to the natural globe. The implant would be porous, smooth and pliable, permitting simple muscle attachment, i.e. without the need for wrapping. It should be cost effective and entirely alloplastic, eliminating potential infectious risks of donated tissue. The implant should incite no inflammation (i.e. should be inert) and be easily distinguished from surrounding tissue on imaging.

The currently available porous implants have a stiff structure (poor compliance). A more compliant implant may be more forgiving of repetitive conjuntival trauma at the implant–ocular prosthesis junction, reducing the incidence of tissue breakdown and implant exposure. Potential candidate

17.10 The original impression of the socket may be used to produce a clear prosthesis with the exact shape, surface irregularities and frictional properties of the patient’s prosthesis. This allows both

the socket and the implant to be observed during movement. The patient illustrated here is a 12-year-old boy who sustained a

severe right-sided ocular injury at age 5 years. At the time of these photographs, he was 4 weeks post-enucleation, with implantation of an unwrapped 18 mm Medpor® ball. NB black mark on center of prosthesis and tissue marker on conjunctiva, placed such that in the primary position the mark on the prosthesis overlaps the tissue marker on the conjunctiva. (a) 1° position, (b) right gaze, poor implant abduction/poor prosthesis abduction, (c) left gaze, good implant adduction/poor prosthesis adduction, (d) upgaze,

both implant and prosthesis elevate well, (e) downgaze, the implant extorts with infraduction, which is not apparent in the movement of the prosthesis.

materials include porous hydrogels, whose high water content allows more similar physical properties to living tissue (Chalasani et al. 2007). A promising design is the AlphaSphere®; this implant incorporates a spongy anterior hemisphere using poly(2-hydroxyethyl methacrylate) (PHEMA) to

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permit suturing of muscles, and a posterior smooth gel hemisphere, which in theory would optimize rotational movement (Hicks et al. 1999; Hicks et al. 2006).

17.8Sources of further information and advice

Illustrations

Illustrations of the exposed integrated and buried integrated motility implants may be found in: Gougelmann H. P., The evolution of the ocular motility implant, International Ophthalmology Clinics, 10, 689, 1970. Original illustrations by Lee Allen, outlining development of the Allen implant may be found in: A buried muscle cone implant – development of a tunneled hemispherical type, Archives Ophthalmology, 43, 879, 1950.

Organizations and websites of interest

∑Amercian Academy of Ophthalmology: www. AAO.org.

∑American Society of Ocularists: www.ocularist.org.

∑American Association for Pediatric Ophthalmology and Strabismus: www.AAPOS.org.

17.9References

Allen, L. 1950, “A buried muscle cone implant – development of a tunneled hemispherical type”, Arch. Ophthalmol., 43, 879–890.

Allen, L. 1970, “Fitting the ocular prosthesis: a challenge”, Trans. Am. Acad.Ophthalmol. Otolaryngol., vol. 74, no. 6, pp. 1318–1320.

Allen, L. 1983, “The argument against imbricating the rectus muscles over spherical orbital implants after enucleation”, Ophthalmology, 90, no. 9, 1116–1120.

Allen, L., Ferguson, E. C., & Braley, A. E. 1960, “A quasi-integrated buried muscle cone implant with good motility and advantages for prosthetic fitting”, Trans. Am. Acad. Ophthalmol. Otolaryngol., 64, 272–286.

Allen, L. & Webster, H. E. 1969, “Modified impression method of artificial eye fitting”,

Am. J. Ophthalmol., 67, 189–218.

Allen. L., Spivey, B. E., & Burns, C. 1969, “A larger Iowa implant”, Am. J. Ophthalmol., 68, 397–400.

Ameniya, T., Matsumura, M., & Hirose, Y. 1977, “Effects of radiation after enucleation without implantation on orbital development of patients with retinoblastoma”,

Ophthalmologica, 174, no. 3, 137–144.

Anderson, R. L., Thiese, S. M., Nerad, J. A., Jordan, D. R., Tse, D., & Allen, L. 1990, “The universal orbital implant: indications and methods”, Adv. Ophthalmic Plast. Reconstr. Surg., 8, 88–99.

Anderson, R. L., Yen, M. T., Lucci, L. M., & Caruso, R. T. 2002, “The quasi-integrated porous polyethylene orbital implant”, Ophthal. Plast. Reconstr. Surg., 18, no. 1, 50–55.